Method of determining depth of compressions during cardio-pulmonary resuscitation

ABSTRACT

A method of processing a raw acceleration signal, measured by an accelerometer-based compression monitor, to produce an accurate and precise estimated actual depth of chest compressions. The raw acceleration signal is filtered during integration and then a moving average of past starting points estimates the actual current starting point. An estimated actual peak of the compression is then determined in a similar fashion. The estimated actual starting point is subtracted from the estimated actual peak to calculate the estimated actual depth of chest compressions. In addition, one or more reference sensors (such as an ECG noise sensor) may be used to help establish the starting points of compressions. The reference sensors may be used, either alone or in combination with other signal processing techniques, to enhance the accuracy and precision of the estimated actual depth of compressions.

FIELD OF THE INVENTIONS

[0001] The methods and devices described below relate to the field ofcardio-pulmonary resuscitation (CPR).

BACKGROUND OF THE INVENTIONS

[0002] The American Heart Association guidelines for the correctapplication of cardio-pulmonary resuscitation (CPR) specify that chestcompressions be performed at the rate of 80 to 100 per minute and at adepth, relative to the spine, of 1.5 to 2.0 inches. (Guidelines 2000 forCardiopulmonary Resuscitation and Emergency Cardiovascular Care, 102Circulation Supp. I (2000).) However, CPR is physically and emotionallychallenging, even for trained professionals. Research has shown thatmanual chest compressions rarely meet the guidelines. See, for example,Ochoa et al., The Effect of Rescuer Fatigue on the Quality of ChestCompressions, Resuscitation, vol. 37, p.149-52. See also Hightower etal., Decay in Quality of Closed-Chest Compressions over Time, Ann EmergMed, 26(3):300-333, Sept. 1995. One of the difficulties of performingcorrect chest compressions is that the rescuer imprecisely judges thetiming and depth of compressions, particularly when the rescuer becomestired. Thus, if accurate and timely user feedback could be provided tothe rescuer then the rescuer would be more likely to perform CPRcorrectly.

[0003] Various devices have been proposed to assist a rescuer inproperly applying CPR. For example, Kelley, Apparatus for Assisting inthe Application of Cardiopulmonary Resuscitation, U.S. Pat. No.5,496,257 (Mar. 5, 1996) shows a device that uses a pressure sensor tomonitor compression forces and timing. Groenke et al., AED with ForceSensor, U.S. Pat. No. 6,125,299 (Sep. 26, 2000) shows a device that usesa force sensor to measure the compression force applied to a patient'schest. However, these devices only measure the force applied to thechest and do not measure the actual depth of compressions. A given forcecan compress the chests of different patients by different amounts, someasuring only force will not provide sufficient or consistent feedbackto the rescuer. In addition, force-based measurements may also beinaccurate because of intra-patient variation in thoracic morphology andcompliance (stiffness).

[0004] CPR devices that use only accelerometers to measure depth ofcompressions, other than our own patented device shown in Halperin etal., CPR Chest Compression Monitor, U.S. Pat. No. 6,390,996 (May 21,2002), do not fully or accurately account for errors in the measuredacceleration; nor do they account for drift in the starting points ofcompressions. In addition, the integration process necessary to derivethe depth of compressions greatly compounds any errors in the measuredacceleration.

[0005] It is important to correct for errors in the measuredacceleration since the total depth of compressions should be within therelatively narrow range of 1.5 inches to 2.0 inches. Numericalsimulations have shown that a total error in acceleration as small as0.02 in/sec² results in an error of 0.25 inches in displacement. Giventhe narrow depth range of optimal compressions, an error of 0.25 inchesis unacceptable. For example, Freeman, Integrated Resuscitation, U.S.Publication 2001/0047140 (Nov. 29, 2001) shows a device that uses anaccelerometer as a compression sensor and mentions gauging chest depthwith the accelerometer. However, Freeman enables no method to accountfor the errors inherent in using an accelerometer alone. Thus anymeasurement Freeman makes of chest compression depth is inaccurate.

[0006] Myklebust et al., System for Measuring and Using ParametersDuring Chest Compression in a Life-Saving Situation or a PracticeSituation and Also Application Thereof, U.S. Pat. No. 6,306,107 (Oct.23, 2001) describes a device which uses a pressure pad, containing anaccelerometer and a force activated switch, to determine the depth ofdepressions. However, Myklebust does not provide a means to measurecompression depth using an accelerometer alone, nor does Myklebustaccount for some kinds of error in the measured value of chestcompression depth (such as drift).

[0007] The problems inherent in the above devices show the difficulty ofsolving the problem of measuring chest compression depth using only anaccelerometer. Nevertheless, the basic concept of determiningdisplacement from a measured acceleration is straightforward (in asystem with a known starting position). Displacement is determined bydouble integrating the measured acceleration.

[0008] However, this method of measuring chest compression depth iscomplicated by at least three major sources of error: signal error,external acceleration error, and drift in the actual or measuredstarting points of compressions from the initial starting point ofcompressions. Signal error comprises errors in the measured accelerationdue to electronic noise, the shaking of wires or cables, errors inherentin the accelerometer, and other sources of noise in the accelerationitself.

[0009] External acceleration error comprises errors introduced byaccelerations applied to the patient and/or the accelerometer other thanaccelerations caused by CPR. For example, if the patient is beingtransported in an ambulance and a rescuer is applying manual CPR with acompression monitor, then the accelerometer will measure accelerationscaused by road vibrations as well as accelerations caused by CPR. (Ifthe ambulance hits a pot hole then a large spike may appear in thecompression waveform.) The accelerometer, by itself, cannot distinguishbetween the accelerations caused by road noise and the accelerationscaused by compressions. In other words, the accelerometer measures acombined acceleration and not just the accelerations caused bycompressions. Accordingly, the compression monitor will report adisplacement value different from the actual chest displacement.

[0010] Another source of error, drift, comprises systematic shifts inthe actual or reported starting points of each compression over anentire series of compressions. The accelerometer has no “memory” of theinitial starting position. Thus, as the rescuer applies compressions thereported depth waveform can start to drift. The compression monitor mayindicate that the reported depth waveform is increasingly deeper thanthe actual waveform. This form of drift is referred to as positivedrift. On the other hand, drift can also cause the compression monitorto report a depth waveform that is increasingly more shallow than theactual waveform. In other words, actual compression starting points arebecoming increasingly deeper, but the compression monitor insteadreports each starting point as close to the initial starting point. Thisform of drift is referred to as negative drift.

[0011] One cause of negative drift is a failure to allow the chest toreturn to a fully relaxed position. Absent correction, the accelerometerwill begin measuring displacement from the new “initial” position. Thus,the compression monitor erroneously informs the rescuer that the currentstarting point is at the initial starting point. However, the actualdepth of the current starting point is more than the depth reported bythe compression monitor. As a result, the rescuer may compress the chestharder than he should to achieve the erroneous depth suggested by thecompression monitor.

[0012] Another source of both types of drift is a change in the overallposition of the accelerometer with respect to the patient. For example,if the accelerometer is not fully secured then the accelerometer maysystematically slip. (This may also cause external acceleration error.)Yet another source of drift is expansion and contraction of the chestdue to ventilation performed simultaneously with compressions. Othersources of drift may also exist. Each source of drift may be independentof the others and may not cancel each other out, so the compressionmonitor should be able to account for both positive and negative drift.

[0013] Notwithstanding drift resulting from erroneous operation, changesin the actual starting point of compressions do occur. For example, ifone or more ribs break during CPR then the actual starting point of eachcompression may be closer to the spine (a phenomena known as chestremodeling). Other types of chest injury or disease that affect thestructure and strength of the rib cage can also cause chest remodeling.Chest remodeling can be gradual, in which case a gradual shift occurs inthe actual initial starting point of compressions. A compression monitorshould be able to account for the difference between erroneous drift andan actual shift in the starting points of compressions.

[0014] These and other sources of error are compounded by integratingthe acceleration. The errors caused by signal noise and drift cause theconstants of integration to have a value other than zero. The non-zeroconstants of integration compound the errors already present in theacceleration. Thus, the total compression depth reported by thecompression monitor can be very inaccurate. Accordingly, methods areneeded to accurately and precisely derive the depth of chestcompressions from a measured acceleration.

SUMMARY

[0015] The methods and devices described below provide for signalprocessing techniques that precisely and accurately derive the depth ofchest compressions from a measured acceleration of chest compressions.Specifically, the methods and devices provided below provide for a meansto correct chest displacement errors caused by signal error, externalacceleration error, and drift. According to one method, a moving averagetechnique is used to produce an accurate measurement of compressiondepth. According to a second method, a change in the patient's ECG(electrocardiogram) may be used to determine the starting points ofcompressions. These methods may be combined together to further increasethe accuracy of chest depth measurement.

[0016] In broad terms, a moving average technique averages a pluralityof compression cycles together, but weights recent compressions moreheavily than compressions further in the past. One moving averagetechnique begins with filtering a raw acceleration signal to eliminateas much signal noise as practicable. The filtered acceleration signal isthen integrated to derive the velocity of compressions. The velocity isfiltered to remove accumulated low frequency variations. The filteredvelocity measurement is integrated again to derive chest displacement.Chest displacement is then processed through a baseline limiter and apeak limiter; the baseline limiter may comprise a moving averageprocessor and the peak limiter may comprise a moving average processor.The baseline limiter estimates the actual starting point of the currentcompression and the peak limiter estimates the actual peak depth of thecurrent compression. A baseline detector then identifies the startingpoint of the current compression. A peak detector then identifies thepeak depth of the current compression. A means for combining signalsthen combines the estimated starting point and the estimated peak depthto derive the estimated actual depth of the current compression.Finally, the estimated actual depth of the current compression isprovided to one or more devices which provide intelligible feedback to amanual CPR provider, to an automated CPR device, or to an ECG operator.

[0017] In another method, a change in the noise component of thepatient's ECG is correlated to the start of a chest compression. Whenthe noise component of the patient's ECG signal exceeds a pre-determinedthreshold then the accelerometer begins to measure acceleration. Thus,the actual starting point of the current compression is established.This method reduces some forms of external acceleration error andreduced drift. The method also helps to set the constants of integrationto zero.

BRIEF DESCRIPTION OF THE DRAWINGS

[0018]FIG. 1 shows a patient and an accelerometer-based compressionmonitor in place on a patient.

[0019]FIG. 2 shows a graph of compression depth over time before signalprocessing, where compression depth is derived from a measuredacceleration.

[0020]FIG. 3 shows a graph of compression velocity over time beforesignal processing, where compression velocity is derived from a measuredacceleration.

[0021]FIG. 4 shows a graph of compression acceleration over time beforesignal processing, where compression acceleration is measured by anaccelerometer.

[0022]FIG. 5 is a flow chart of a signal processing technique thatconverts a raw compression acceleration into an estimated actualcompression depth.

[0023]FIG. 6 is a flow chart of an alternate signal processing techniquethat converts a raw compression acceleration into an estimated actualcompression depth.

[0024]FIG. 7 shows the graph of compression depth over time afterfiltering the raw acceleration.

[0025]FIG. 8 shows the graph of compression velocity over time afterfiltering the raw acceleration.

[0026]FIG. 9 shows the graph of compression acceleration over time afterfiltering the raw acceleration.

[0027]FIG. 10 shows the graph of compression depth over time afterfiltering both the raw acceleration and the derived velocity.

[0028]FIG. 11 shows the graph of compression velocity over time afterfiltering both the raw acceleration and the derived velocity.

[0029]FIG. 12 shows the graph of compression acceleration over timeafter filtering the raw acceleration.

[0030]FIG. 13 shows the graph of compression depth over time afterfiltering both the raw acceleration and the derived velocity, and afterapplying a baseline limiter to the compression depth waveform.

[0031]FIG. 14 shows the graph of compression velocity over time afterfiltering both the raw acceleration and the derived velocity, and afterapplying the baseline limiter to the compression velocity waveform.

[0032]FIG. 15 shows the graph of compression acceleration over timeafter filtering the raw acceleration and after applying the baselinelimiter to the compression acceleration waveform.

[0033]FIG. 16 shows the graph of compression depth over time afterfiltering both the raw acceleration and the derived velocity, and afterapplying the baseline limiter and the peak limiter to the compressiondepth waveform.

[0034]FIG. 17 shows the graph of compression velocity over time afterfiltering both the raw acceleration and the derived velocity, and afterapplying the baseline limiter and the peak limiter to the compressionvelocity waveform.

[0035]FIG. 18 shows the graph of compression acceleration over timeafter filtering the raw acceleration and after applying the baselinelimiter and the peak limiter to the compression acceleration waveform.

[0036]FIG. 19 is a flow chart of a signal processing technique that usesa change in ECG noise to activate a switch which, in turn, controls whenan accelerometer begins to measure acceleration.

[0037]FIG. 20 shows a graph of compression depth over time before signalprocessing and with a negative drift in the reported compression depthwaveform.

[0038]FIG. 21 shows a graph of compression velocity over time beforesignal processing and with a negative drift in the reported compressionvelocity waveform.

[0039]FIG. 22 shows a graph of compression acceleration over time beforesignal processing and with a negative drift in the reported compressionacceleration waveform.

[0040]FIG. 23 shows the graph of FIG. 20 corrected by using a change inECG noise to establish the actual starting points of compressions.

[0041]FIG. 24 shows the graph of FIG. 21 corrected by using a change inECG noise to establish the actual starting points of compressions.

[0042]FIG. 25 shows the graph of FIG. 22 corrected by using a change inECG noise to establish the actual starting points of compressions.

[0043]FIG. 26 shows an accelerometer-based compression monitor in placeon a patient and a system of reference sensors comprising a referenceaccelerometer, a switch, and a load sensor disposed such that eachsensor may measure various parameters related to chest compressions.

[0044]FIG. 27 illustrates a compression waveform that a user feedbacksystem may prompt the rescuer to perform.

[0045]FIG. 28 is a block diagram of how an actual chest compressionacceleration is converted into a corrupted value for chest position.

[0046]FIG. 29 is a block diagram of a general solution for converting acorrupted chest compression acceleration into an estimated actual depthof chest compressions.

[0047]FIG. 30 is a block diagram of how an actual ECG signal isconverted into a corrupted ECG signal.

[0048]FIG. 31 is a block diagram of a general solution for converting amotion corrupted ECG signal into an estimated actual ECG signal.

[0049]FIG. 32 is a graph of a pig's ECG signal that is corrupted bynoise caused by chest compressions.

[0050]FIG. 33 is a graph of CPR motion where CPR is performed on a pig.

[0051]FIG. 34 is a graph of the pig's estimated ECG noise signal.

[0052]FIG. 35 is a graph of the pig's estimated actual ECG signal.

DETAILED DESCRIPTION OF THE INVENTIONS

[0053]FIG. 1 shows a patient 1 and an accelerometer-based compressionmonitor 2 in place on the patient. An accelerometer-based compressionmonitor uses one or more accelerometers to determine the depth ofcompressions. An example of an accelerometer-based compression monitormay be found in our own patent, Halperin et al., CPR Chest CompressionMonitor, U.S. Pat. No. 6,390,996 (May 21, 2002), which is herebyincorporated by reference in its entirety. The compression monitor 2 isplaced on the sternum 3 of the patient 1, on the rescuer's hands orarms, or on an automatic CPR device. The chest is then compressed. Theaccelerometer measures the acceleration of compressions and a processor4 estimates the actual displacement of the accelerometer based on themeasured acceleration. The signal processing techniques described belowensure that the estimated actual displacement is accurate and precise.

[0054] The estimated actual displacement may be provided to adisplacement display 5 that provides intelligible feedback to a manualCPR provider or to an automated CPR device. Likewise, other CPR-relatedparameters may be provided to one or more compression device displays 6(or other means for user feedback). CPR-related parameters include thedepth of chest compressions, the velocity of chest compressions, theacceleration of chest compressions, and the patient's ECG.

[0055] In the case of the patient's ECG, the compression monitor may beprovided with one or more electrodes. The processor may process thepatient's ECG during compressions to produce an estimated actual ECG.The estimated actual ECG may then be provided to an ECG display 7 (orother means for user feedback) that provides intelligible feedback tothe manual CPR provider, to an automated CPR device, or to otherindividuals or devices that monitor the patient's ECG.

[0056] The following terms are used throughout the specification and aredefined as follows:

[0057] Actual compression depth: the actual depth of a compression atany given time.

[0058] Actual starting point of a compression: the actual place or pointat which a chest compression begins.

[0059] Autoregressive moving average: a function that uses past datasamples to modify the current data sample.

[0060] Baseline portion of the compression depth waveform: that portionof depth waveform where the set of actual starting points is most likelyto be found.

[0061] Baseline limiter: a processor or function that operates on thebaseline portion of the compression depth waveform.

[0062] Compression Peak: the place or point where maximum compressiondepth occurs.

[0063] Current compression depth: the depth of a compression at anygiven time.

[0064] Current starting points: the starting point of the currentcompression.

[0065] Depth of compressions: the depth the chest is compressed at anyinstant in time, where depth is measured relative to the relaxedposition of the chest.

[0066] Estimated actual starting point of a compression: the estimatedvalue of the actual place or point at which a chest compression begins.

[0067] Initial starting point of compressions: the place or point atwhich a series of compressions begins.

[0068] Measured starting point of a compression: the measured value ofthe place or point at which a chest compression begins.

[0069] Moving average: a function that uses past data samples to modifythe current data sample.

[0070] Past starting points: the starting points of compressions thathave already occurred.

[0071] Peak portion of the compression depth waveform: that portion ofdepth waveform where the set of actual peaks are most likely to befound.

[0072] Starting point of a compression: the place or point at which achest compression is begun.

[0073]FIGS. 2 through 4 show graphs of compression depth, velocity, andacceleration over time for four hypothetical compressions. No signalprocessing has been applied to any of waveforms shown in FIGS. 2 through4. Compression depth in FIG. 2 is shown as a positive value-the higherthe value, the deeper the chest has been compressed. The phantomwaveforms 12 represent the actual waveforms for compression depth,velocity, and acceleration (measured independently of theaccelerometer). The solid waveforms 13 represents the waveforms derivedfrom the acceleration measured by the compression monitor accelerometer.The waveforms 13 are also the waveforms reported by the compressionmonitor to the signal processing system 4. Compression depth is measuredin inches, marked at 1 inch intervals, compression velocity is measuredin inches per second (in/s), marked at 1 in/s intervals, and compressionacceleration is measured in inches per second per second (in/s²), markedat 1 in/s² intervals. For all three Figures time is measured in seconds,marked at 1 second intervals. The start of compressions is at time equalto zero. The initial depth of compressions is at depth equal to zero.

[0074] Phantom lines 14 and 15 intersect all three graphs. Phantom line14 corresponds to the time at which maximum compression depth isobtained. Phantom line 15 corresponds to the time at which minimumcompression depth is obtained. In addition, phantom line 14 indicatesthat a compression depth maximum 16 corresponds to a compressionvelocity of zero. Phantom line 14 also indicates that an accelerationmaximum 17 is slightly offset from the compression depth maximum 16.Likewise, phantom line 15 indicates that a compression minimum 18 (orstarting point or zero point) corresponds to a compression velocity ofzero. Phantom line 15 also indicates that an acceleration minimum 19 isslightly offset from the compression depth minimum 18. A compressionvelocity maximum 20 and minimum 21 occur around the middle of acompression.

[0075] The solid waveforms show the effects of three major types oferror: signal error, external acceleration error, and drift. Signalerror is primarily represented by the “noisy” (rough) nature of thesolid waveforms; however, external acceleration error can also form aportion of the “noise.” Although the acceleration waveform is lessnoisy, integrating the acceleration increases the effect of the noise inthe velocity waveform. Integrating the velocity waveform increases theeffect of the noise yet again. Thus, the compression depth noise FIG. 2is higher than the compression velocity noise in FIG. 3, which is inturn higher than the compression acceleration noise in FIG. 4.Accordingly, the compression monitor will report a very noisycompression depth waveform.

[0076] External acceleration error is primarily represented by thelarge, positive spike 22 in the solid waveforms of FIGS. 2 through 4.(Although the spike in FIGS. 2 through 4 occurs at a maximum, spikes canoccur anywhere in the compression cycle and can affect the measuredacceleration both positively and negatively). The spike is caused by alarge acceleration unrelated to compressions, but nevertheless measuredby the accelerometer. Thus, the actual waveform 12 in all three figuresshows a corresponding peak 23 significantly below spike 22. Accordingly,absent the correction suggested here, the compression monitor willreport for that compression cycle a compression depth much higher thanthe actual compression depth.

[0077] Drift is primarily represented by the increasing distance betweenthe respective minimums of the actual and reported waveforms of FIGS. 2through 4, as shown by arrows 24 and 25. The drift is causing thecompression monitor to erroneously report a compression waveform that isbecoming increasingly deeper (positive drift). However, the actualwaveform is more closely returning to the initial starting point, and isthus the drift shown in FIGS. 2 through 4 is considered a positivedrift. Likewise, arrows 24 and 25 in FIGS. 3 and 4 illustrate that drifthas an increasing affect on the reported velocity and the reportedacceleration. The effects of drift mean that the initial starting pointof compressions cannot be used as a reliable starting point for allcompressions. Accordingly, the starting point of compressions must bedetermined for every compression cycle. In addition, the other sourcesof noise must be either eliminated or greatly reduced.

[0078]FIG. 5 is a flow chart of a signal processing technique thatconverts a raw acceleration into an estimated actual value for totalcompression depth. The raw acceleration 34 is filtered by a first filterin step 35 to produce a filtered acceleration. The first filtercomprises a high-pass filter and greatly reduces most forms of signalnoise. (In other embodiments the first filter may comprise a band passfilter, a moving average filter, an infinite impulse response filter, anautoregressive filter, or an autoregressive moving average filter.) Theeffects of the other steps shown in FIG. 5 are described in the contextof FIGS. 7 through 18.

[0079]FIG. 6 is a flow chart of an alternate signal processing techniquethat converts a raw compression acceleration into an estimated actualcompression depth. This flowchart is described after the description forFIGS. 7 through 18.

[0080] The effect of the filter operation 35 is seen in FIGS. 7 through9, which show the graphs of compression depth, velocity, andacceleration over time for four hypothetical compressions after thefirst filtering step 35. (FIGS. 7 through 9 show the output of the firstfiltering step). The measured acceleration waveform 13 of FIG. 9 is muchless noisy than the corresponding unfiltered waveform 13 of FIG. 4.Since the velocity and depth waveforms of FIGS. 8 and 9 are derived fromthe acceleration waveform they, too, are less noisy. Nevertheless, theintegration process still causes the velocity waveform to be more noisythan the acceleration waveform and the depth waveform to be more noisythan the velocity waveform. In addition, the external acceleration spike22 still remains, as do the errors caused by drift (as shown by arrows24 and 25).

[0081] Returning to FIG. 5, the filtered acceleration is integrated in afirst integration step 36 to derive the compression velocity. However,as shown in FIG. 8, without further processing the velocity waveform isstill noisy. Thus, the velocity is filtered by a second filter in step37 to produce a filtered velocity. The second filter comprises a highpass filter and further reduces most signal noise in the velocity anddepth waveforms. (In other embodiments the second filter may comprise aband pass filter, a moving average filter, an infinite impulse responsefilter, an autoregressive filter, or an autoregressive moving averagefilter.)

[0082] The effects of the filter operation 37 is seen in FIGS. 10through 12, which show the graphs of compression depth, velocity, andacceleration over time for four hypothetical compressions after thesecond filtering step 37. (FIGS. 10 through 12 show the output of thesecond filtering step 37.) The measured velocity waveform 13 of FIG. 11is less noisy than that of FIG. 8 (the velocity waveform after the firstfiltering step). Since the depth waveform is derived from the velocitywaveform it, too, is correspondingly less noisy. Nevertheless, theintegration process still causes the depth waveform to be slightly morenoisy than the acceleration and velocity waveforms. In addition, theexternal acceleration spike 22 still remains, as do the errors caused bydrift (as shown by arrows 24 and 25).

[0083] Returning to FIG. 5, the filtered velocity is integrated in asecond integration step 38 to calculate the chest compression depth.Signal noise has been substantially eliminated and thus a thirdfiltering step is not required. However, the noise in the depthwaveform, as shown in FIG. 10, is still slightly more than the noise inthe velocity waveform, as shown in FIG. 11. Thus in other embodiments athird filter, comprising a high pass, bandpass, or other filter may beused to further reduce signal noise in the depth waveform.

[0084] After the initial filtering steps (35 and 37) and integrationsteps (36 and 38), a baseline limiter estimates the actual startingpoint of a compression in step 39. The baseline limiter uses, amongother techniques described below, the starting points from pastcompressions to estimate the current compression starting point. Thebaseline limiter itself comprises a digital or analog signal processorthat operates on the baseline portion of the compression depth waveformof FIG. 10. The baseline portion of the compression depth waveformcomprises that portion of depth waveform where the set of actualstarting points is most likely to be found. For example, the baselinemay comprise the portion of the depth waveform that is equal to andbelow 1.1 inches compression depth. (Larger changes in the startingpoints of compressions are unlikely, and signals indicating largechanges are probably wrong.) Thus, the limiter will disregard orarbitrarily assign a realistic depth value to any “starting point” above1.1 inches depth. In one embodiment, past starting points above thebaseline are disregarded and a current starting point above the baselineis reported or treated as an error. (Past starting points are thestarting points of compressions that have already occurred. A currentstarting point is the starting point of the current compression.) Inanother embodiment a current starting point above the baseline isassigned a small probability and averaged with the past starting points.

[0085] In one embodiment the baseline limiter estimates the startingpoint of the current compression by applying a moving average to allstarting points that fall within the baseline portion of the depthwaveform. A moving average is a function that uses past data samples tomodify the current data sample. (Additional moving average techniquesare described below.) In the case of the baseline limiter, the baselinelimiter may weigh recent starting points more heavily than olderstarting points, meaning that the weight of a given starting pointdecays over time. Starting points that fall outside the baseline portionof the depth waveform are given an arbitrary weight or no weight. Byapplying a moving average to all starting points the baseline limiterreduces the effect of external acceleration error and drift on thecurrent starting point. In other words, the moving average of allstarting points will be statistically closer to the current actualstarting point than the current measured starting point derived from theintegration of the acceleration.

[0086] The following example shows an embodiment of a moving averagetechnique. In this embodiment each compression starting point is given aweight of 1.25% of the previous compression starting point. In otherembodiments the weighting may comprise a percentage in the range ofabout 0.1% to about 12.5% (which yields between about 0.3% to about 90%data weighting at the end of about 1 minute). In other words, themeasured value of the current starting point (starting point 1) isweighted 100%, the most recent starting point (starting point 2) isweighted 98.75%, the next previous starting point (starting point 3) isweighted 97.5%, the next previous starting point (starting point 4) isweighted 96.25%, etc until all compressions are weighted. Eventually,compressions in the distant past are given no virtually no weight atall. The depth of all the weighted starting points is then averaged. Theweighted average of all starting points is treated or reported as thecurrent starting point.

[0087] In another embodiment, all compressions after a pre-determinedtime period (such as about 1 minute to about 15 minutes) aredisregarded. Thus, only compressions within the last 1 to 15 minutes areaveraged. In another embodiment, all compressions after a pre-determinednumber of compressions (such as about 5 to about 15) are disregarded.

[0088] Continuing the example, in one embodiment the measured values forstarting point 1=0.5 inches, starting point 21.1 inches, starting point3=4.0 inches, and starting point 4=0.9 inches. Starting point 3 isoutside the baseline portion of the depth waveform (the baseline portionis 1.1 inches and below in this example). Starting points outside thebaseline in this example are disregarded, so starting point 3 isdisregarded. Thus, the current starting point would be reported as:

[(0.5*100%)+(1.1*98.75%)+(0.9*96.25%)]÷3=0.853 inches

[0089] relative to the initial starting point.

[0090] Had starting point 3 been included in the moving average, thenthe current starting point would have been reported as:

[(0.5*100%)+(1.1*98.75%)+(4.0*97.5%)+(0.9*96.25%)]÷4=1.615

[0091] inches relative to the initial starting point.

[0092] Stated differently, this value is the estimated actual startingpoint for the current compression.

[0093] Mathematically, the reported value of the current starting pointis expressed as:

Ds=[Σ(DB _(i)*ω^(i−1))]÷n_(r)

[0094] where DB_(i)=0 if DB_(i)>B,

[0095] where Ds is the depth of the current starting point, n_(r) is thenumber of starting points remaining after all starting points thatexceed the baseline have been disregarded, i is the starting pointnumber (or sum index), DBi is the measured depth of the i^(th) startingpoint, ω is the weighting constant, and B is the baseline. Expresseddifferently, DB_(i)*ω^(i−1) is summed from i=1 to n_(r) and the sum isdivided by n_(r), but if a particular DB_(i) is greater than B then thatDB_(i) is instead set to zero.

[0096] The baseline limiter may perform other functions to furtherincrease the accuracy and precision of the estimated depth of thecurrent starting point. For example, a probability can be assigned to agiven change between the current starting point and the immediateprevious starting point. (Likewise a probability can be assigned to agiven change between the current starting point and the moving averageof all previous starting points.) Large changes in starting point may begiven less weight than smaller changes. This technique may be referredto as a “weighted moving average” technique.

[0097] Continuing the above example, measured depth 1 is treated ashaving a 100% probability of occurring. Then, the difference between thecurrent starting point (depth 1) and the previous starting point (depth2) is 1.1 inches −0.5 inches=0.6 inches. The probability of a step of0.6 inches occurring is assigned to be 97%, based on past experiments.Since the probability is not 100%, the current starting point is nottreated as having jumped a full 0.6 inches. Instead, the currentstarting point is treated as having jumped 0.6*0.97 0.582 inches.Accordingly when calculating the weighted moving average depth 2 istreated as being 1.082 inches and not 1.1 inches. Starting point 3 isstill disregarded. The difference between starting point 2 (1.1 inches)and starting point 4 (0.9 inches) is 0.2 inches, which is assigned a 99%probability. Thus, the effective distance of the step between depth 2and depth 4 is 0.2*99%=0.198. Accordingly, depth 4 is treated as 0.902inches instead of 0.9 inches. Using the same moving average as above,the current starting point is now reported as:

[(0.5*100%)+(1.082*98.75%)+(0.902*96.25%)]÷3=0.812 inches

[0098] relative to the initial starting point.

[0099] Stated differently, this value is the estimated actual startingpoint for the current compression.

[0100] Mathematically, the reported value of the current starting pointis expressed as:

Ds={Σ[DB _(i−j)+(DB _(i) −DB _(i−j))*P _(S)]*ω^(i−1) ]}÷n _(r)

[0101] where DB_(i)=0 if DB_(l)>B,

[0102] where Ds is the depth of the current starting point, n_(r) is thenumber of starting points remaining after all starting points thatexceed the baseline have been disregarded, i is the starting pointnumber, DBi is the measured depth of the i_(th) starting point, j is theindex for the most recent starting point that was still within thebaseline, DB_(i−j) is the most recent starting point that was stillwithin baseline, P_(s) is the probability that a step of sizeDB_(i)−DB_(l−j) will occur, ω is the weighting constant, and B is thebaseline. The result, Ds, is the reported depth of the current startingpoint. Expressed differently, [DB_(i−j) 30(DB_(i)−DB_(l−j))*P_(s)]*ω^(i−1) is summed from i=1 to n and the sum isdivided by n_(r), but if a particular DB_(l) is greater than B (thebaseline) then that DB_(i) is instead set to zero.

[0103] In another embodiment, a probability is assigned to the step sizebetween the depth of the current starting point and the weighted averageof all previous starting points. (In the above example, the probabilityis assigned to a step size between the current starting point and theimmediate past starting point). This technique may be referred to as a“weighted moving average with memory” technique. In this technique thereported depth of the current starting point is expressed mathematicallyas:

Ds={Σ[DB _(l−j)+(DB _(i) −Ds _(i−j))*P _(S)]*ω^(l−1) ]}÷n _(r)

[0104] where Ds_(1−j)=[Σ(DB _(i−j)*ω^(i−2))]÷n_(r) and DB_(l)=0 ifDB_(l)>B,

[0105] where the variables are defined above. Again, the value for Ds isalso the estimated actual starting point for the current compression.

[0106] In another embodiment, an autoregressive moving average (ARMA)filter may be used as the baseline limiter. The ARMA filter is anexponentially decaying “forgetting” filter that weights more currentdata more heavily than past data. The ARMA operates on more than justthe compression starting point or peak values. Instead, the ARMA filteroperates on data samples of compression acceleration, velocity, or depthtaken at rapid time intervals. Data samples may be taken at a rate ofabout 100 samples per second to about 2000 samples per second (with arate of about 1000 samples per second preferred). Thus, the ARMA filteroperates on the entire waveform and not just on the compression peaksand the starting points.

[0107] In low pass form (which eliminates high frequency variations inthe baseline) the ARMA filter may be expressed mathematically as:

y[n]=(1−α)*y[n−1]+α*x[n].

[0108] In this case, n is the index of the current sample (the “n^(th)”sample), y[n] is the output of the current sample, x[n] is the input ofthe current sample, y[n−1] is the output from the previous sample, and αis an independent term that determines how fast the filter “forgets”past outputs and the amount of influence the current input has on theoutput. The value for α may be in the range of about 0.02 to about0.0002, with a value of about 0.002 being suitable for many CPR-relatedfilter applications. Should it be desired to implement a high-pass ARMAfilter for the baseline limiter, then the ARMA equation becomes:

y[n](high pass)=1−{(1−α)*y[n−1]+α*x[n]},

[0109] where y[n](high pass) is the high pass filter output and theother variables are defined in the context of the low pass ARMA filter.The high pass filter may be used to eliminate low-frequency variationsin the depth, velocity, or acceleration signals.

[0110] The moving average techniques in the above examples have beendescribed in the context of processing the compression depth waveform.However, the techniques can be used to process the velocity waveform andthe acceleration waveform, should it be desired to report accuratevalues for the velocity and acceleration of compressions. The movingaverage techniques may be applied to each waveform separately. In otherwords, one does not necessarily apply a moving average technique to theacceleration waveform, then integrate the acceleration waveform, thenapply a second moving average technique to the velocity waveform, thenintegrate the velocity waveform, and finally apply a third movingaverage technique to the depth waveform. However, in other embodimentsthis procedure may be used.

[0111] Other methods for analyzing the baseline signal may be used todetermine the estimated actual starting point of compressions. Anotherembodiment of the baseline limiter comprises a signal processor thatuses a transition probability map to identify the probability ofparticular shifts in the measured starting point. (The probability mapmay be pre-determined, such as by using a density estimator or kernelestimator, and then hard-coded into the compression monitor software.) Aparticular starting point measurement is compared to the probability mapand the system determines by how much a given shift in the measuredstarting point is erroneous. The reported starting point is adjustedaccordingly. (Likewise, a transition probability map may be used toestimate the actual peak and also the actual maximum depth for eachcompression.) The effect of the baseline limiter 39 is seen in FIGS. 13through 15, which show the graphs of compression depth, velocity, andacceleration over time for four hypothetical compressions. FIGS. 13through 15 also show the output of steps 35 through 39 in FIG. 5. Thebaseline limiter has been applied separately to the velocity waveform(FIG. 14) in step 47 and to the acceleration waveform (FIG. 15) in step48.

[0112]FIGS. 13 through 15 show that a moving average technique reducesthe effect of drift in the reported starting point of each compression.(The moving average techniques also reduce the effect of externalacceleration errors that appear in the baseline portion of thewaveform). Before correction, the reported starting points were becomingincreasingly deeper, though the actual starting points were returning toclose to the actual initial starting point. By applying a moving averagetechnique to the baseline of a measured waveform, the reported startingpoints of each compression are statistically closer to the actualstarting points. Accordingly, the compression monitor will report anestimated actual compression depth that is closer to the actualcompression depth. Arrows 49 and 50, which are shorter than arrows 24and 25 in FIGS. 2 through 4 and FIGS. 7 through 12, show the beneficialeffect of applying a moving average technique to each waveform.

[0113] Returning to FIG. 5, the compression depth waveform corrected bythe baseline limiter may be passed through a third filter in step 51 toreduce any accumulated signal noise in the compression depth waveform.The third filter comprises a high pass filter, though in otherembodiments the third filter may comprise a band pass filter.

[0114] Subsequently, the depth waveform (whether filtered or unfiltered)is provided to a starting point detector in step 52. The starting pointdetector identifies the value of the current estimated starting point.The current estimated starting point is then provided to a means forcombining signals 53 (as indicated by line 54). The means for combiningsignals 53 will later use the current estimated starting point tocalculate the estimated actual compression depth. The means forcombining signals comprises a signal adder, a linear system model, anon-linear system model, or other means for combining signals.

[0115] Next, the compression waveform may be provided to a peak limiterin step 55. The peak limiter is a signal processor that performs similarfunctions to the baseline limiter, but instead operates on the peakportion of a compression waveform. The peak portion of the waveformcomprises that portion of the waveform in which a peak is most likely tooccur. In one embodiment, the peak portion is the portion of thewaveform above the baseline portion. Continuing the example given forthe baseline limiter, the peak portion of the depth waveform would bethe portion of the depth waveform that is above 1.1 inches. The peaklimiter thus will smooth the peak portion of a waveform in much the sameway as the baseline limiter smoothes the baseline portion of a waveform.

[0116] In one embodiment the peak limiter sets an outside boundary onthe size of the maximum compression depth. Thus, the peak limiter eitherdisregards (throws out) or sets an arbitrary value to any peak that isgreater than a known, improbable peak value (the depth of a largeperson's chest, for example, would not be a probable value for CPRcompression depth). Thus, the peak limiter prevents the compressionmonitor from reporting a compression depth that is improbable.

[0117] The effect of the peak limiter is seen in FIGS. 16 through 18,which show the graphs of compression depth, velocity, and accelerationover time for four hypothetical compressions after the peak limiter step55 in FIG. 5. (FIGS. 16 through 18 show the output of steps 35 through55). A peak limiter has been applied separately to the velocity waveformin step 56 and to the acceleration waveform in step 57. By applying amoving average technique to the peak portion of the compressionwaveforms, the effect of the external acceleration spike 22 has beengreatly reduced. Combined with the techniques discussed in the previousprocessing steps, the reported waveforms are now close to the actualwaveforms.

[0118] Returning to FIG. 5, the estimated peak may optionally beprovided to a fourth filter 58 to remove remaining signal noise. Thefourth filter comprises a high pass filter, though in other embodimentsthe fourth filter may comprise a band pass or other filter.

[0119] Subsequently, the depth waveform is provided to a peak detectorin step 59. The peak detector identifies the value of the estimated peak(the estimated maximum depth of the current compression). The estimatedpeak is then provided to the means for combining signals 53. The meansfor combining signals 53 combines the estimated starting point 52 withthe estimated peak 59 to produce an estimated actual compression depthfor the current compression 61. The estimated actual depth is thenprovided to a means for user feedback 62 (a user feedback system). Themeans for user feedback may comprise a speaker, a visual display, one ormore LEDs, a vibrator, radio, or other means for communicating with therescuer. The user feedback system in turn provides informationcorresponding to the estimated actual depth of the current compressionto the rescuer.

[0120] In the technique of FIG. 5, the baseline portion and the peakportion do not overlap. Thus, the compression depth waveform may bethought of as comprising two portions, the baseline portion and the peakportion. Each portion of the depth waveform is treated differently bytwo different procedures (the baseline limiter and the peak limiter) toextract different information. Thus, both the baseline limiter and thepeak limiter operate on the same depth waveform. The effect of this isthat the signal comprising the depth waveform is provided first to thebaseline limiter and then to the peak limiter (the signal is not split).

[0121] The technique shown in FIG. 6 may be used when the baselineportion and the peak portion overlap (though the technique may also beused when the baseline portion and peak portion do not overlap). Forexample, the technique of FIG. 6 may be used when the baseline portionis set below 1.5 inches (relative to the chest's relaxed position) andthe peak portion is set above 1.0 inches (relative to the chest'srelaxed position). In this case the signal representing the depthwaveform is split and is provided to two separate processors, a baselinelimiter and a peak limiter. Each processor performs similar functions tothe limiters already described. Thus, although the baseline limiter andthe peak limiter act independently of each other, the technique of FIG.6 produces an estimated starting point and an estimated peak in much thesame was as the technique shown in FIG. 5. The means for combiningsignals then combines the estimated starting point and estimated peak instep 53 to produce the estimated actual depth of the currentcompression. The estimated actual depth of the current compression isprovided to the user feedback system in step 62. The user feedbacksystem in turn provides the estimated actual depth of the currentcompression to the rescuer.

[0122] In addition to the signal processing techniques of FIGS. 5 and 6,other techniques can be used to correct for errors in the compressiondepth waveform. For example, FIG. 19 is a flow chart of a signalprocessing technique that uses a change in ECG noise 63 to activate aswitch 64 that, in turn, controls when an accelerometer begins tomeasure acceleration.

[0123] To implement this technique, the compression monitor is providedwith one or more electrodes, or some other means for measuring thepatient's ECG. As the rescuer performs compressions the patient's ECGbecomes noisy. Even if the patient's actual ECG is flat (shows noactivity) the reported ECG will still show the noise caused by chestcompressions. Indeed, a motion artifact signal (an ECG noise componentcaused by chest compressions) will be superimposed on any ECG rhythm.Whatever the actual ECG rhythm, the ECG noise may be isolated andaccounted for.

[0124] Since the bulk of ECG noise during compressions is caused by theact of compressing the chest, the starting point of a compression may becorrelated to the point where the ECG noise exceeds a pre-determinedthreshold. However, there is some delay or lag between the onset of acompression and the onset of ECG noise. The time lag is on the order ofmilliseconds to tenths of a second. In order not to miss any part of acompression, a buffer (either digital or analog) may be employed tocorrect for the time lag. Thereafter, when the ECG noise exceeds theparticular threshold then the switch is programmed to activate theaccelerometer (which will begin to take acceleration measurements).Total compression depth is then determined by double integrating themeasured acceleration.

[0125] The effect of using ECG noise as a reference sensor to establishthe starting points of compressions is seen in FIGS. 20 through 25,which show compression depth, velocity, and acceleration over time forfour hypothetical compressions. No signal processing is applied to anyof waveforms shown in FIGS. 20 through 22. The phantom waveforms 12represent the actual waveforms for compression depth, velocity, andacceleration (measured independently of the accelerometer). The solidwaveforms 13 represent the waveforms derived from the accelerationmeasured by the accelerometer. The solid waveforms are also thewaveforms reported by the compression monitor. The effects of signalnoise are shown by the rough nature of the solid waveforms. The effectsof external acceleration noise are shown by the two spikes, 65 and 66,in the reported waveform. The effects of negative drift (increasinglyshallow compressions) are shown by the increasing distance (representedby arrows 67 and 68) between the minimums in the reported and the actualwaveforms.

[0126] The effect of using ECG noise as a reference sensor to establishthe starting points of compressions is seen in FIGS. 23 through 25,which show graphs of compression depth, velocity, and acceleration overtime for hypothetical compressions. Using ECG noise as a referencesensor reduces certain external acceleration errors and reduces theeffect of negative drift. (The ECG noise reference sensor can alsoreduce the effect of positive drift). Specifically, the ECG noisereference sensor reduces the effect of external acceleration noise thatoccurs near a compression minimum. Since the accelerometer is not “on,”a portion of the external acceleration spike is “ignored”. In practicethe accelerometer is still taking data, but software or hardware is usedto process out accelerometer data or signals that occur during a timeperiod where ECG noise does not reach a predetermined level. In othermethods, the estimated actual depth of compressions is calculated whenthe ECG noise falls within a predetermined threshold. In any case, theeffect of spike 65 is reduced in the reported waveform. However, theaccelerometer by itself still cannot tell the difference between acompression-related acceleration and an external acceleration. Thus, thereported waveform is still subject to external acceleration noise thatoccurs during a compression, as shown by spike 66.

[0127] Nevertheless, the ECG noise reference sensor does reduce theeffects of drift. Since the starting point of a compression isindependently established, the waveform is much less subject to eitherpositive or negative drift. In other words, the accelerometer willalways measure acceleration after the actual start of compressions.Thus, the reported waveform of FIG. 23 more accurately shows what therescuer is actually doing compressing the chest from starting pointsthat are becoming increasingly deep. Thus, peaks 69 and 70 show that themeasured waveform more closely matches the actual waveform.

[0128] Although the ECG noise reference sensor can reduce the effects ofdrift and reduce the effect of some forms of external accelerationnoise, signal noise remains a problem. Thus, FIGS. 23 through 25 stillshow the same levels of signal noise as shown in FIGS. 20 through 22. Toreduce all forms of noise the ECG noise reference sensor may be combinedwith the signal processing techniques of FIG. 5 or 6. The combinedtechniques will produce a reported depth waveform that is close to theactual waveform.

[0129] Other reference sensors may be used to establish the actualstarting point of a compression. FIG. 26 shows an accelerometer-basedcompression monitor in place on a patient 1 who is lying on a surface80. A system of reference sensors comprising an accelerometer 81, a loadsensor 82, and a switch 83 are disposed such that each sensor maymeasure various parameters related to chest compressions. In the case ofreference accelerometers, the reference accelerometers may be disposedelsewhere on the patient, or upon any reference object that experiencesthe same external accelerations the patient experiences. The referenceaccelerometers may comprise a three-axis accelerometer, but may alsocomprise three orthogonal single-axis accelerometers or one single axisaccelerometer (in which case the accelerations along the other two axesare assumed to be negligible).

[0130] The reference accelerometers 81 allow a signal processor toeliminate external acceleration error, such as those accelerationscaused by transporting the patient. In one method, the accelerationsensed by the compression monitor or automatic CPR device (the deviceacceleration) is provided to a signal processor. The device accelerationcontains the acceleration caused by compressions (the compressionacceleration) and the acceleration caused by the external accelerations(the external acceleration). Next, the reference accelerometer oraccelerometers provide a reference acceleration to the signal processor.The reference acceleration contains only the external acceleration ofthe patient. Then the reference acceleration is combined with the deviceacceleration to produce an estimated actual acceleration. (The effect ofcompression accelerations on the reference acceleration is negligiblesince the surface and patient are kept steady with respect to thecompression monitor.)

[0131] Once obtained, the estimated actual acceleration may be doubleintegrated to produce an estimated actual chest depth. Thus, the depthof compressions may be determined even in the presence of large externalaccelerations. Moreover, the position signal may be made more accurateand precise by combining the actual acceleration with the signalprocessing technique of FIG. 5 or 6, or with other signal processingtechniques.

[0132] In lieu of (or in addition to) the ECG noise sensor and referenceaccelerometers, other reference sensors may be used to set the actualstarting point of a compression. Reference sensors may comprise a loadsensor 82, a switch 83, a transthoracic impedance detector, an ECG noisedetector (as described above), a voltage or current sensor in anautomatic CPR device, a start signal in an automatic CPR device, anencoder in an automatic CPR device, or any other sensor capable ofindependently detecting the actual beginning of a compression. When thereference sensor detects the beginning of a compression then thestarting point is set to zero. The acceleration is then processed toderive compression depth. The technique of setting the starting point tozero when a reference sensor detects the beginning of a compression mayalso be combined with the signal processing techniques of FIG. 5 or 6.

[0133] In the case of a switch 83, the switch is disposed such that whena compression begins the switch will be closed. For example, the switchmay be disposed beneath or on the compression monitor, on the patient 1,on the surface 80 upon which the patient lies, on the rescuer's hand, ona CPR machine, on the patient, or on some other location that allows theswitch to register that a compression has begun.

[0134] The switch may comprise many different types of switches andsensors, including a contact switch, a motion sensor, a voltage sensoron an automatic CPR device, an optical, rotary, or other encoder on anautomatic CPR device, the displacement of a shaft or other component onan automatic CPR device, a potentiometer, a strain gage, apiezoresistive transducer, a differential transformer, synchro andinduction potentiometers, variable-inductance and variable-reluctancepickups, an eddy current non-conducting transducer, a capacitivetransducer, an electro-optical transducer, a photographic switch, avideo tape switch, a holographic switch, a switch that uses photoelastictechniques, translation encoders, an ultrasonic transducer, moving coiland moving magnet pickups, an AC or DC tachometer, an eddy-currentdrag-cup tachometer, additional accelerometers, or a gyroscopicdisplacement switch.

[0135] In the case of the load sensor 82, the load sensor may beoperatively connected to the rescuer, the patient, an automatic CPRdevice, beneath the patient, or elsewhere so long as the load sensorsenses a load when compressions begin. When the load sensor measures aload that exceeds a pre-determined threshold, then the measured startingpoint is set to zero. The load sensor may also be operatively connectedto a switch, which activates when the load sensor senses a load, or theload sensor may merely provide input to a signal processor systemidentifier (described in more detail below). Compression depth is thendetermined by integrating the acceleration twice. The technique ofsetting the starting point to zero when a load sensor detects thebeginning of a compression may also be combined with the signalprocessing techniques of FIG. 5 or 6.

[0136] In another embodiment of the load sensor 82, the load sensor maybe disposed such that the sensor can sense both the weight of thepatient and the force of compressions. The load sensor 82 may bedisposed beneath the surface 80 upon which the patient 1 rests. Duringcompressions the force of pressing on the patient causes the load sensorto report a total force greater than the patient's weight. Accordingly,a starting point is set to zero when the total force is about equal tothe patient's weight.

[0137] Examples of force sensors that can be used with this techniqueinclude pressure sensors, elastic force transducers, shaft displacementon an automatic CPR device, a voltage or a current sensor on anautomatic CPR device, an optical, rotary, or other encoder on anautomatic CPR device, bonded strain gages, beam strain gages,differential transformers, piezoelectric transducers, variablereluctance/FM oscillators, gyroscopic force transducers, and vibratingwire force detectors. Examples of pressure sensors that can be used withthis technique include deadweight gages, manometers, elastictransducers, piezoelectric transducers, and force-balance transducers.

[0138] In the case of a transthoracic impedance detector, one or moreECG, defibrillation, or other electrodes are disposed on the patient'sthorax. When a compression begins the impedance of the thorax changes.The thoracic impedance comprises the impedance due to skin and thoraciccontents between any two electrodes. The change in thoracic impedancemay be measured by a small test current or by any other means formeasuring impedance. When the impedance changes by a pre-determinedamount then the starting point is set to zero. Total compression depthmay then be determined by processing the measured acceleration.

[0139] Because the compression monitor can measure the compressionwaveform, the compression monitor can also prompt the rescuer or anautomatic CPR device to perform a particular compression waveform. FIG.27 shows a compression waveform that the compression monitor may promptthe rescuer to perform. Depth is measured in inches and time is measuredin seconds. The scale shown in FIG. 27 is marked in 0.5 second intervalsand 1.0 inch intervals respectively. The compression phase of the cycleis indicated by the positively sloped curve 84. The compression phase ofthe cycle ends at the maximum compression depth 85 (compression peak).The decompression phase of the cycle is indicated by the negativelysloped curve 86. The decompression phase ends when the rescuer begins anew compression at the next starting point 87 (or baseline), which mayor may not be at the initial starting point. Compressions are initiatedat time=0 and depth=0, and the total depth of compressions is thedistance represented by arrows 88.

[0140] The compression waveform includes a compression hold 89, wherethe rescuer maintains a hold at maximum compression depth for a shortperiod of time, and an incomplete decompression hold 90, where therescuer maintains a short hold at a point deeper than the initialstarting point. Each compression and decompression is performed quickly,at high acceleration and velocity, as indicated by the relatively steepslopes of the compression phase 84 and the decompression phase 86. Theduty cycle is slightly less than 50% (the ratio of compression anddecompression time is 1), meaning that slightly less time is spent inthe compression phase, as indicated by the distance between arrows 89,than in the decompression phase, as indicated by the distance betweenarrows 90.

[0141] Although the compression waveform of FIG. 27 shows an example ofa particular waveform that the compression monitor can instruct arescuer to perform, other waveforms are also possible. For example,another waveform may lack a compression hold phase. Yet another variesthe duty cycle and others increase the compression hold time. The exactwaveform depends on the current state of the art of what kind ofcompression waveform comprises an optimal compression waveform for aparticular kind of patient. In addition, the compression monitor may beprovided with a switch, button, software, or other means for user inputwhich allows the rescuer to enter the size or shape of the patient. Thecompression monitor may use this information to choose a particularwaveform from a library of waveforms. The compression waveforms are thusadaptable to findings in future research, AHA guidelines, rescuerobservations, and medical professional preferences. Accordingly, atvarious times different waveforms may be provided to the user feedbacksystem, as described more fully below.

[0142] The prompted waveform may be provided by the user feedback system(step 62 in FIG. 5). In addition, the user feedback system may providethe rescuer or automatic CPR device with other compression-relatedinformation. For example, the user feedback system may displayinformation regarding the starting point of compressions, thecompression depth waveform, the compression velocity waveform, and thecompression acceleration waveform. Thus, the user feedback system mayprovide the rescuer or the automatic CPR device with all of the dataneeded to continuously track the position, velocity, and acceleration ofthe chest during all phases of CPR. This information may be used toevaluate the performance of a rescuer or automatic CPR device.

[0143] The user feedback system may also provide a rescuer or automaticCPR device with information concerning the compression phase quality andthe decompression phase quality. Compression phase quality is thequality of compressions with respect to total compression depth, theduty cycle, the acceleration of compressions, smoothness ofcompressions, and other factors related to the compression phase.Decompression phase quality is the quality of compressions with respectto whether the rescuer returns to the actual initial position, the dutycycle, the acceleration of decompressions, the smoothness ofdecompressions, and other factors related to the decompression phase.The rescuer or automatic CPR device may use this information to evaluateor prompt the kind and quality of compressions.

[0144] The user feedback system 62 may provide the rescuer or automaticCPR device with information concerning compression phase quality bycombining information gained from the acceleration, velocity, andposition waveforms. For example, the user feedback system can instructthe rescuer to increase compression force when the depth of compressionsare less than recommended guidelines and to reduce compression forcewhen the depth of compressions are greater than recommended guidelines.The user feedback system may also instruct the rescuer with regard toother compression phase parameters of a compression waveform. Forexample, the user feedback system can inform the rescuer or automaticCPR device if the time to achieve proper compression depth is too shortor too long.

[0145] The user feedback system 62 may also provide the rescuer orautomatic CPR device with information regarding decompression phasequality by combining information gained from the acceleration, velocity,and position waveforms. For example, the user feedback system caninstruct the user or device on the proper position at which to restafter a decompression. Thus, the feedback system can instruct the useror device to allow the chest to fully relax if the rescuer or device isnot allowing the chest to fully return to its initial starting position.Conversely, should it be medically indicated, the user feedback systemcan instruct the user or the device to return to a depth just below theinitial chest position. In this case, the rescuer or device implements a“decompression hold” and maintains force on the chest even when thecompression cycle reaches its minimum depth. In another case thefeedback system can indicate different compression starting points atdifferent times. Thus, the user feedback system can instruct the rescueror device to apply incomplete decompression holds during compressioncycles, but to allow the chest to return to its fully relaxed positionduring ventilation pauses. The user feedback system may also instructthe rescuer or device with regard to other decompression phaseparameters, such as the decompression rate and the duty cycle of thedecompression phase.

[0146] Taken together, the information gained from the compression phasequality and the decompression phase quality enable the user feedbacksystem to prompt the rescuer on how to perform an optimum compressionwaveform and an optimum compression duty cycle. A rescuer performs aparticular compression waveform by performing compressions at apre-determined depth and rate, and by holding the chest at apre-determined compression depth for a pre-determined time. A rescuerperforms a particular duty cycle by compressing the chest for apre-determined period and allowing the chest to relax for anotherpre-determined period.

[0147] Thus, the user feedback system can prompt the rescuer orautomatic CPR device to perform at the appropriate compression rate,compression depth, compression velocity (the time required to compressor decompress the patient), compression acceleration, and compressionhold time for each phase (compression and decompression) of thecompression cycle. Accordingly, the compression waveform that therescuer or device actually applies can conform to a complex compressionwaveform. Since research has shown that most patients benefit from morecomplex waveforms, patient survival is likely to increase if the rescueror automatic CPR device uses a compression monitor with this userfeedback system.

[0148] Similarly, the user feedback system 62 of FIG. 5 can provide therescuer or CPR device with feedback regarding the compression dutycycle. The duty cycle is the ratio of time under compression to the timeunder decompression for each compression cycle. (However, the duty cycledoes not include time periods where no compressions are taking place,such as during ventilation.) If the duty cycle does not fall withinpre-determined parameters, then the user feedback system may prompt therescuer to adjust compression timing and compression rate in order toeffect an optimal duty cycle.

[0149] The user feedback system described above comprises the last stepin a particular solution to the problem of determining an accurate valuefor chest displacement from a raw acceleration signal. (FIG. 5 is theflowchart for this solution). Many variations of that solution exist, asalready described, though it is possible to view the problem from ageneral perspective and to create a general solution.

[0150]FIGS. 28 and 29 are block diagrams that represent the generalproblem to and the general solution for determining an accurate positionfrom an acceleration measured during CPR. FIG. 28 is a block diagram ofhow an actual chest compression acceleration is converted into acorrupted value for chest position. In broad terms, the actualacceleration 105, signal noise 106, external acceleration noise 107, andsome forms of drift 108 are combined by an unknown function 109 (whichmay be linear or non-linear and may include random or deterministicinputs). The unknown non-linear function is known as the system, whichproduces the corrupted acceleration 110 measured by the accelerometer.The corrupted acceleration is then integrated twice, which greatlycompounds the problem introduced by the corruption in the acceleration.The increased error is referred to as integration error 111 (although itis assumed that the integration technique itself does not directlycontribute errors into the position). Finally, additional sources ofdrift 112 can affect the final value for the corrupted position 113.

[0151]FIG. 29 is a block diagram of a general solution for converting acorrupted chest compression acceleration into an estimated actual depthof chest compressions. First, a reference sensor 119 may establish theactual starting point of a compression. Thus, the starting point of theacceleration will be known. (Although helpful, the reference sensor 119is not necessary to the general solution). The actual acceleration 105and the real or the estimated noise sources 120 (which comprise blocks106 through 108 of FIG. 28) are combined by the system 109 by an unknownfunction. The result is the corrupted acceleration 110. The measuredacceleration is then provided to a means for combining data 121 (whichmay comprise a linear or a non-linear function) and to a systemidentifier 122.

[0152] The system identifier comprises one or more functions (eitherlinear or non-linear) that model the system. One or more noisereferences that can be correlated to the noise sources 120 may also beprovided to the system identification function 122. For example, noiseidentified by a low frequency filter can be correlated to signal noiseor a reference accelerometer can be correlated to the externalacceleration noise.

[0153] The system identification function may also use variousparameters of an automatic CPR device as noise source references, evenif the reference itself does not produce noise in the acceleration.However, the noise source reference must somehow be correlated to asource of noise in the acceleration signal. For example, theaccelerometer-based depth measurement reports a chest depth of 0.5inches. However, a simultaneous current spike in the automatic CPRdevice informs the system that the CPR device is compressing the chestmuch harder than should be required to achieve a chest depth of 0.5inches. The discrepancy may be caused by external acceleration noise orby drift. Thus, the current spike may be correlated to a source of noisein the system. This information may be used by the system identifier tohelp model the system. Likewise, voltage, shaft displacement, or opticalor rotary encoders may be used as references by the system identifier tohelp model the system. (Again, the noise references are useful but notnecessary).

[0154] The system identifier then combines or correlates the noisesource references and the measured acceleration in order to produce theestimated noise 123 in the measured acceleration. The estimated noise123 is then provided to the means for combining data 121. The means forcombining data combines the estimated noise 123 and the measuredacceleration 110 to produce an estimated actual acceleration 124. Theestimated actual acceleration is then integrated 125 twice. Filters 126may optionally be used during one or both integration steps to reducethe compounding effect of errors that may still linger in the estimatedactual acceleration. The final result is an accurate and preciseestimate of the actual position 127 of the accelerometer.

[0155] The system identification function 122 models the system and thuscan be used to estimate the noise in the acceleration. (Once the noiseis known it can be easily eliminated by combining the noisy accelerationwith the measured acceleration.) In other words, system identificationis the process of using the input and output data to model the functionthat combines the actual acceleration and the sources of noise in theacceleration. The system identification problem has a known or measuredoutput and an input that may be known or unknown. The addition of knownor measured input is beneficial to system identification, but notnecessary. The system itself is an unknown arbitrary function that canbe linear or non-linear, though some boundary conditions may be known.

[0156] A number of methods, both linear and non-linear, may be used tomodel the system. Each of these methods may comprise, alone or incombination, the system identification function in step 122. Thesemethods may operate by taking many data samples per second, as opposedto operating only on the compression starting points or peak points.Nevertheless, these methods may also be performed on the compressionstarting points or the compression peaks. A partial list of thesemethods include: autoregression, autoregression with extra inputs,autoregressive moving average (which is one of the methods used in thetechniques shown in FIGS. 5 and 6) autoregressive moving average withextra inputs, autoregressive integrated moving average, autoregressiveintegrated moving average with extra inputs, a Box-Jenkins model, anoutput error model, a hidden Markov model, a Fourier transform, awavelet transform, wavelet de-noising, wavelet filtering, adaptiveneural networks, recurrent neural networks, radial basis function nets,adaptive curve fitting (splines), Kalman filters, extended Kalmanfilters, adaptive Kalman filters, unscented Kalman filters, and kernelestimation. Algorithmic approaches that may be used to find the systemidentification function include maximum entropy, maximum likelihood,recursive least squares (or similar techniques), numerical methods,unconstrained global search or optimization, expectation minimization,and fast Fourier transforms.

[0157] In the case of recursive identification, the formula for generalrecursive identification may be expressed as:

X(t)=H(t, X(t−1), y(t), u(t)) and  (1)

θ(t)=h(X(t)),  (2)

[0158] where X(t) is the state of the system at time t; H is the stateof the transfer function; X(t−1) is the past system state; Y(t) is themeasured output; u(t) is the measured input; θ(t) is the system, and htransforms the system state to the output. The system state can beconverted to the system output by h(X(t)).

[0159] Since X(t) and θ(t) are evaluated at each time point as u(t) andy(t) are collected, the total amount of previously collected data has amuch more profound effect on the system than do the most recentlycollected data.

[0160] Equations (1) and (2) may be simplified into equations (3) and(4):

X(t)=X(t−1)+μQ_(x)(X(t−1), y(t), u(t)) and  (3)

θ(t)=θ(t−1)+γQ_(θ)(X(t−1), y(t), u(t)),  (4)

[0161] where μ and γ are small numbers that reflect the relative amountof information provided by the latest time step. Q is a function thatrelates inputs, outputs, and states. Equations (3) and (4) are moresimple in the sense that it is more simple to compute the equations byrecursive than equations (1) and (2).

[0162] A number of numerical algorithms may be used to solve equations(3) and (4). A partial list of numerical algorithms include recursiveleast squares, recursive (or recursion) instrumental variables,recursive prediction error methods, recursive pseudolinear regression,recursive Kalman filters (including time varying parameters), andrecursive Kalman filters for time varying systems. These numericaltechniques encompass many of the famous “named” techniques as specialcases, including a Kalman filter, an extended Kalman filter, extendedrecursive least squares, and others. Each algorithm has strengths andweaknesses, but all asymptotically approach a solution to equations (3)and (4).

[0163] The “named” special cases may be derived from general equations(3) and (4) when certain conditions or assumptions are made. Thus, theequations for each of the listed algorithms can be further specified.For example, when using a recursive least squares algorithm equation (4)may be expressed as:

θ(t)=θ(t−1)+L(t)[y(t)−φ^(T)(t)θ(t−1)] where  (5) $\begin{matrix}{{L(t)} = {\frac{{P\left( {t - 1} \right)}{\varphi (t)}}{{\lambda (t)} + {{\varphi^{T}(t)}{P\left( {t - 1} \right)}{\varphi (t)}}}\quad {and}}} & \left( {5a} \right) \\{{P(t)} = {\frac{P\left( {t - 1} \right)}{\lambda (t)} - {\frac{{P\left( {t - 1} \right)}{\varphi (t)}{\varphi^{T}(t)}{P\left( {t - 1} \right)}}{{\lambda (t)}\left\lbrack {{\lambda (t)} + {{\varphi^{T}(t)}{P\left( {t - 1} \right)}{\varphi (t)}}} \right.}.}}} & \left( {5b} \right)\end{matrix}$

[0164] In equations 5 through 5(b) L(t), P(t), and P(t−1) are terms usedto simplify the equation, φ(t) is a regression vector, λ(t) is aforgetting factor (described in more detail below), and φ^(T)(t) is thetranspose of the regression vector.

[0165] In addition, equation 4 can also be expressed for the cases ofrecursive instrumental variables, recursive prediction-error methods,recursive pseudolinear regression, a recursive Kalman filter fortime-varying systems, and a recursive Kalman filter with parametricvariation.

[0166] Once the system identification algorithm as been selected fromthe above set of algorithms, there are several additional parametersthat may affect the quality of the model. These additional parametersinclude data weighting, choice of updating step, choice of updatinggain, and model order selection. In the case of data weighting, when asystem is time-varying the input-output data near the present time moreaccurately reflects the nature of the present system. Data recordedfurther back in time is more closely related to a past system state. Toreflect this fact the data can be weighted to favor a more recent systemstate. Actual data weighting is accomplished by the “forgetting factor,”λ, in equations 5 through 5b. The selection of λ is made based oninformation about how fast the system changes state. A typical range forλ is between about 0.9800 and about 0.9999 (though λ may be 1.0000 if no“forgetting factor” is desired).

[0167] Another way of thinking about the effect of λ on the systemidentifier is to evaluate at what point a data sample is given a weightof about 36%. (36% is the value of the number e⁻¹, which is the value atwhich a data sample may be considered statistically insignificant). Atthis weight a data sample's statistical significance becomes relativelysmall. The sample number at which a data sample has a weight of about36%, known as T₀, may be mathematically expressed as:$T_{0} = {\frac{1}{1 - \lambda}.}$

[0168] T₀ (and hence λ) is selected with appropriate knowledge of thesystem state and can be used to tune system identification so that theestimated actual acceleration most closely approximates the actualacceleration if the actual acceleration were independently measured.Thus, T₀ is pre-set before the compression monitor begins takingmeasurements.

[0169] The closer λ is to 1 the more samples are needed to reach thepoint where a given data sample is given a weight of about 36%. Asmaller λ means that a given data sample is “forgotten” more quickly.For example, if λ is equal to 0.9800 then T₀=50 samples, meaning thatthe 50^(th) sample receives a weight of about 36%. However, if λ0.9999then T₀=10,000, meaning that sample number 10,000 receives a weight of36%. In the upper limit, if λ=1 then T₀=∞, meaning that a data sample isnever “forgotten” (it always receives a weight of 100%).

[0170] The sampling rate (how many times a second the acceleration ismeasured) affects the how λ changes the system identifier. If samplesare taken 1000 times per second then data may be “forgotten” rapidly onthe time scale of CPR compressions. For example, if the sample rate is1000 times per second and T₀=1000 then data from just 1 second in thepast is given a weight of 36%. In practice, the sampling rate may varyfrom about 100 samples per second to about 2000 samples per second. Auseful sample rate for signal processing acceleration measurementsduring CPR is about 500 samples per second. In other embodiments thesample rate may be faster, but every certain number of samples may beignored. For example, samples may be taken at 1000 samples per second,but every other sample ignored. This process, known as decimation, hasthe same effect as a slower sample rate.

[0171] In the preceding discussion the forgetting factor was a fixednumber; it did not change with time. However, λ can vary with time sothat the system identifier may adapt to changing situations. Forexample, λ may vary during a ventilation pause and in one embodiment λincreases during a ventilation pause. The effect of increasing λ duringventilation pauses is to discard data points very quickly. Thus, thecompression monitor will not report a change in compression depth duringa ventilation pause.

[0172] In addition to adding a forgetting factor to the systemidentification function, the choice of updating step affects the qualityof the model. (Although only some of the system identificationtechniques require an update; for example, a Kalman filter requires anupdating step). The update step can be implemented using a variety ofmethods. Some system identifiers may be solved analytically, such as theKalman filter, and the updating step may be solved analytically. Othersystem identifiers must be solved numerically. Three updating methodsthat may be used when a numerical solution is required are aGauss-Newton update, a gradient update, and a Levenberg-Marquardtupdate. The Gauss-Newton update converges to an accurate fit of theactual solution, though it requires a large number of steps (and thusmore computation time). The gradient update converges quickly but doesnot converge as accurately to the actual solution as the Gauss-Newtonupdate. The methods may be combined. The gradient update is used firstto converge the fit quickly and then the identifier switches to theGauss-Newton update to achieve the final fit. This combined technique isknown as a Levenberg-Marquardt update.

[0173] Mathematically the Gauss-Newton update may be expressed as:

R(t)=R(t=1)+γ(t)[φ(t)φ^(T)(t)−R(t−1)].

[0174] Mathematically the gradient update may be expressed as:

R(t)=I|φ(t)|² =R(t−a)+γ(t)[I|φ(t)|² −R(t−1)].

[0175] In both equations R(t) is the Hessian of the identificationcriterion, R(t−1) is the Hessian of the identification criterion in theprevious time step, γ(t) is the updating gain (which is related to theforgetting factor), and φ(t) is the regression vector.

[0176] The choice of updating gain is another step that is used in manyrecursive system identification functions. The choice of updating gainmay be expressed mathematically as:

y(t)=(1+λ(t)/γ(t−1))⁻¹.

[0177] Thus, the updating gain is related to the forgetting factor.

[0178] With regard to model order selection, the recursive systemidentification techniques fit a system model to input-output data. Thestructure of that model must be determined before the recursive. Thestandard way of solving the model structure problem is to solve a widerange of model structures and then select which model fits the databest. A simple measure of model fit, like the mean squared error, tendsto over-estimate the model order and fit to the process or measurementnoise. If a model is under-estimated, critical components of the systemmight be missed. Several measures of model fit, or metrics, areevaluated over a range of model orders. The model order is the number ofterms used in the model. The smallest model order that minimizes the fitis the appropriate model.

[0179] Several techniques can be used to estimate smallest model order,including final prediction error (FPE), Akaike's information criterion(AIC), maximum description length (a variant of the AIC), andstatistical hypothesis testing (such as the student's t-test). The finalprediction error can be expressed mathematically as: $\begin{matrix}{{{FPE} = \frac{V\left( {1 + {d/N}} \right)}{\left( {1 - {d/N}} \right)}},} & (11)\end{matrix}$

[0180] where V is the quadratic loss function, d is the size of themodel order, and N is the number of data points.

[0181] Akaike's information criterion may be expressed as:

AIC=log[V(1+2(d/N))],

[0182] where V is the quadratic loss function. The quadratic lossfunction may be any quadratic function that relates the additional costfunction of using additional terms.

[0183] The system identification techniques described above have beendescribed in the context of solving the problem of estimating actualcompression depth from a raw acceleration measurement. These techniquesmay also be used to process a noisy ECG signal. FIG. 30 is a blockdiagram illustrating the problem of ECG noise caused by CPR and othersources of noise. Stated differently, FIG. 30 is a block diagram of howa theoretical actual ECG signal 135 and the noise sources are combinedto produce the measured ECG (which contains the motion corruptingartifact). The ECG noise, comprising ECG noise due to compressions 136and other sources of noise 137, are combined with the actual ECG by anunknown linear or non-linear function known as the system 138. Theprimary source of noise in the ECG is due to the motion of compressions,though other sources of noise exist and may be accounted for by thesolution presented below. The system produces a corrupted ECG 139 which,if left unprocessed, cannot be used to accurately report the electricalactivity of the patient's heart. In addition, the system combines theECG noise and the actual ECG in a way that causes the ECG noise tooverlap the actual ECG in the frequency domain. Thus, a simple bandpassfilter is insufficient to accurately process the corrupted ECG. (Asimple bandpass filter will eliminate important components of the actualECG as well as eliminating the ECG noise).

[0184]FIG. 31 is a block diagram of a general solution to the problemillustrated in FIG. 30 and illustrates the process of converting amotion corrupted ECG signal into an estimated actual ECG signal. As withFIG. 30, the system 138 combines the actual ECG 135 and the ECG noise136 to produce the corrupted ECG 139 that is measured by an observer.Next, the measured ECG 139 and a reference corresponding to the ECGnoise 136 are provided to a system identifier 140. For example, sinceCPR induced motion is the largest cause of ECG noise, a signalcorresponding to CPR induced motion may be provided to the systemidentifier. Specifically, a force transducer may be disposed on acompression monitor (or on the patient or rescuer) such that the forcetransducer measures force during a compression. A signal correspondingto the force is provided to the system identifier as a reference signal.Other signals corresponding to CPR induced motion may comprise variousparameters of an automatic CPR device. For example, a signal correlatingto the displacement of a drive shaft or other component can becorrelated to the CPR motion, a signal corresponding to the change incurrent or voltage required to drive the device can be correlated to theCPR motion, or signals produced by optical or rotary encoders may becorrelated to the CPR motion.

[0185] The system identifier models the system and then estimates thenoise component of the measured ECG signal (the estimated noise 141).The estimated ECG noise 141 and the measured ECG 139 are then providedto a means for combining signals 142, which combines the ECG noise andthe measured ECG to produce the estimated actual ECG 143. Since theestimated actual ECG is produced during compressions, the signalprocessing method allows the ECG sensor to detect the heart's normalsinus rhythm even during compressions. Thus, there is no need toperiodically pause compressions to check for the existence of a pulse.As a result, the overall quality of CPR increases and the patient ismore likely to survive.

[0186] The system identifier 140 may comprise similar kinds of functionsand methods as described in the context of the signal processing methodsof FIG. 29. For example, the recursive least squares method described inthe context of FIG. 29 may be used to identify the noise component ofthe measured ECG signal.

[0187]FIGS. 32 through 35 show the effect of using the method of FIG. 31to estimate a pig's actual ECG when the ECG is measured during chestcompressions. For all four graphs each time marker 150 along time scale151 corresponds to the same time marker in the other three graphs, thusmaking possible a direct comparison of one graph to each of the othergraphs. However, the voltage scales 152 of FIGS. 32, 34, and 35 areslightly different from each other.

[0188]FIG. 32 is a graph (millivolts versus milliseconds) of an actualpig's ECG signal that is corrupted by noise caused by chestcompressions. FIG. 32 represents the ECG measured during compressionswithout signal processing.

[0189]FIG. 33 is a graph of force versus time for an actual CPR motionsignal. The motion signal comprises a time varying force signal andcorresponds to the force a CPR device places on the pig's chest whileperforming chest compressions. The force peaks 153 correspond to themaximum depth of compressions. In the case of a compression monitor, themotion signal could comprise a time varying force signal thatcorresponds to the force placed on the patient's chest while a rescueror automatic CPR device performs chest compressions. In this case aforce transducer disposed on a compression monitor (such as on the backof the compression monitor) measures the force of compressions andproduces the force signal. The force signal is later correlated to theECG noise. The force transducer or other force sensor may also bedisposed under the patient's back and then operably connected to thecompression monitor.

[0190]FIG. 34 is a graph of voltage versus time for the estimated ECGnoise signal caused by the chest compressions shown in FIG. 33. Acomparison of FIGS. 33 and 34 shows that the time varying pressuresignal corresponds directly to incidence of ECG noise. In other words,the pressure peaks 153 caused by chest compressions correspond to theincidence of noise peaks 154.

[0191] The system identifier 140 used to generate the estimated noisecomponent of the noisy ECG comprises a recursive least squares method asdescribed in the context of FIG. 29. The autoregressive order wasselected to be equal to 1. The moving average order was selected to be10. The autoregressive order was selected to be 10. The derivative orderwas also selected to be 0. (The derivative order is a linear ornon-linear term used in the system model; specifically it may be eithera truncated positive derivative or a truncated negative derivative. Thenon-linear terms are extensions of the recursive least squares model fitalgorithm). The forgetting factor, λ, was selected to be 1.0000.

[0192]FIG. 35 is a graph of the pig's estimated actual ECG signal. Thegraph of FIG. 35 is generated by subtracting the estimated noise signalof FIG. 34 from the measured ECG signal of FIG. 32. The estimated actualECG signal corresponds closely to the pig's actual ECG signal.

[0193] The signal processing methods described in the context of noisyECG signals (FIGS. 30 and 31) and noisy acceleration signals (FIGS. 28and 29), as well as the techniques described in relation to the baselinelimiter and the peak limiter, may also be used to estimate the actualvalue of the patient's transthoracic impedance (the chest's electricalresistance or impedance). The estimated actual value of the patient'stransthoracic impedance may be used to determine the amount of voltageneeded to shock the patient with a defibrillator.

[0194] As compressions are applied to the patient the measured value ofthe transthoracic impedance becomes noisy. The general signal processingsolutions and the limiters already described may be used to identify,isolate, and eliminate the noise component of the measured transthoracicimpedance. Thus, the actual value of the transthoracic impedance may beestimated.

[0195] The estimated actual value of the transthoracic impedance may beprovided to a means for defibrillating the patient. The means fordefibrillating the patient uses the estimated actual value of thetransthoracic impedance to determine the exact voltage necessary toapply an effective shock to the patient. Since the exact value of thevoltage required is also known, the defibrillator can be usedefficiently (thus preserving battery life and making the device safer).

[0196] Since both the estimated actual ECG and the estimated actualtransthoracic impedance are known, an automated CPR device equipped withan AED (automated external defibrillator) may perform defibrillationshocks to a patient without stopping compressions. The device maydetermine when defibrillation is appropriate based on the estimatedactual ECG and may apply the appropriate defibrillation voltage based onthe estimated actual transthoracic impedance. Since compressions do notstop during defibrillation, the patient's blood flow does not stop(meaning the patient is more likely to survive).

[0197] The compression monitor using these signal processing techniques(for either chest depth measurement or ECG measurement) may be used withany means for compressing the chest of the patient. A means forcompressing the chest may comprise manual CPR, electro-stimulation, ameans for performing automatic CPR (including belts, straps, pistons,and plates that are driven by motors or manual levers), or other devicessuitable for compressing the chest. Examples of automatic CPR devicesmay be found in our own patent Sherman et al., Modular CPR AssistDevice, U.S. Pat. No. 6,066,106 (May 23, 2000) and in our applicationCPR Assist Device with Pressure Bladder Feedback, U.S. application Ser.No. 09/866,377 filed May 25, 2001. (Both devices use optical or rotaryencoders disposed on a compression belt or a drive shaft or spool tomeasure the amount of belt pay out or compression depth). Theaccelerometer itself may be disposed in any location where theaccelerometer experiences the downward or upward acceleration ofcompressions. For example, the accelerometer may be disposed within orotherwise disposed on the means for compressing the chest, such as acompression belt. (In the case of manual compressions, the compressionmonitor may be disposed beneath a rescuer's hands while the rescuerperforms compressions, or the compression monitor may be otherwisedisposed on the rescuer's hands, wrist, or arms.)

[0198] If the compression monitor is provided with a means for sensingthe tilt of the accelerometer, such as a three-axis accelerometer,three-axis load sensor, three-axis displacement measurement device, orthe tilt sensor shown in our own U.S. Pat. No. 6,390,996 to Halperin etal., then the user feedback system can prompt the rescuer with respectto compression efficiency. For most patients, compressions are mostefficient when compressions are performed perpendicular to the sternum(straight down in most cases). The tilt sensor measures the angle atwhich compressions are performed and the compression monitor prompts therescuer to adjust the angle if the angle falls outside a particularrange.

[0199] The compression monitor and signal processor may also be operablyconnected to a defibrillator. While the rescuer or device is performingcompressions the defibrillator or compression monitor tracks thepatient's ECG. If the compression monitor's processor measures an ECGsignal that indicates that the patient would benefit from a shock, thenthe rescuer would be instructed to apply a defibrillation shock or toallow an AED to administer a shock. A means for estimating the patient'sactual ECG during compressions comprises our own method disclosed inU.S. Pat. No. 6,390,996 to Halperin et al and the method disclosed inthis application.

[0200] The compression monitor may also be operably connected to a meansfor performing ventilation. After the rescuer has performed anappropriate number of compressions, such as 15, the compression monitorwill instruct the rescuer to pause compressions. The means forperforming ventilation will then administer an appropriate number ofventilations. After ventilations the compression monitor may evaluatethe patient's condition. If the patient still requires compressions,then the compression monitor will instruct the rescuer to resumecompressions. The means for performing ventilation may comprise therescuer, a bag or balloon, a positive pressure ventilator, anelectro-ventilator such as those shown in our own patent Sherman et al.,Chest Compression Device with Electro-Stimulation, U.S. Pat. No.6,213,960 (Apr. 10, 2001), or other means for performing ventilation.

[0201] The compression monitor may also be provided with a means forcommunication that allows the compression monitor to communicate with aremote network. The means for communication comprises a signal carrierand a compression monitor communicator. The signal carrier may comprisea telephone line, direct connection cables, a dedicated digital network,a cell phone network, a satellite communication array, radio or otherelectromagnetic waves, a LED, the internet, or other means for carryinga signal. The compression monitor communicator may comprise a radio orother electromagnetic wave transmitter and receiver, a LED, a modem, orother means for transmitting and receiving a signal carrier. The meansfor communication allows the compression monitor to upload or downloadinformation from a remote network. The compression monitor may also beprovided with a global positioning satellite reader (GPS reader),speakers, keypads, telephones, modems, microphones, cameras, or visualdisplays to allow the user to receive and input information. Data may beexchanged by the means for communication using known communicationstandards, such as the bluetooth standard.

[0202] The remote network may provide other information to thecompression monitor and may also receive information from thecompression monitor. For example, the compression monitor may be updatedwith additional waveforms, patient treatment history, ventilationratios, or other compression-related information of use in a subsequentor current emergency. In the case when a group of monitors are assignedto a single network, each compression monitor may be tracked separatelyby the remote network and provided with different information (ifwarranted). The remote network itself may comprise one or morecomputers, the internet, another CPR-related device, or a human operatorcapable of remotely programming the compression monitor or remotelyprompting the rescuer.

[0203] In use, the compression monitor is provided at a location, suchas a shopping mall or a public place, and is stored until needed. Uponactivation, the compression monitor establishes communication between itand the remote network, which may be a computer located in a callcenter. Emergency responders, such as police, fire, and ambulanceservices, may be notified of the activation and directed to go to alocation pinpointed by the GPS reader or to contact the personactivating the compression monitor. The remote network computer and anoperator trained to use the computer may provide voice assistance to therescuer while monitoring real time streaming data from the compressionmonitor. The operator or computer may from time to time provide otherinformation to the compression monitor or rescuer. For example, thecompression monitor may be provided with data corresponding to whatwaveform the rescuer should be prompted to perform and the operator mayverbally coach the rescuer. In turn, the compression monitor providesdata, such as compression depth, rate, force, waveform, and duty cycle,to the operator and computer for medical analysis. Should the rescuerbecome fatigued then the operator can provide a substitute waveform thatis easier for the rescuer to follow. Likewise, the operator can provideverbal encouragement to the rescuer.

[0204] Another use for the means for communication is to enableautomatic or prompted maintenance of the compression monitor. Eitherperiodically or continuously the compression monitor communicates with aremote network and transmits data such as battery life and status,number of uses, whether any parts need to be replaced, and otherinformation concerning maintenance status. In response, the compressionmonitor either automatically performs maintenance on itself, such as asoftware upgrade, or it prompts a user to perform maintenance upon it.Another use for the means for communication is to enable the compressionmonitor to communicate with additional products used during an emergencyresponse. Example of such a products include those products describedabove, a drug dispenser, or any other product useful for responding tothe emergency.

[0205] An example of combined product use begins with a rescuerbeginning manual resuscitation with the compression monitor. Emergencymedical personnel arrive and deploy an automated chest compressiondevice equipped with an AED. The automated chest compression device isadapted to exchange information with the compression monitor. After theautomated chest compression device is deployed, the compression monitorautomatically communicates with the chest compression device andtransfers relevant treatment history to it, such as time undercompression, quality of compressions as compared to an ideal waveform,ECG history, and other relevant medical data. Based on the informationprovided by the compression monitor, the automated chest compressiondevice may provide a defibrillation shock to the patient beforebeginning compressions. Conversely, the automated chest compressiondevice may determine, based on the transferred data, that chestcompressions must be continued before administering a shock.

[0206] Another example of a combined product is a compression monitorand a separate signal processor. The signal processor may be provided asone or more physical chips (hardware) or may be provided as a computerprogram (software). In either case, the signal processor may be aseparate unit or module and need not be built directly into the monitor.Accordingly, the signal processing units may be provided as stand-aloneproducts. Thus, while the preferred embodiments of the devices andmethods have been described in reference to the environment in whichthey were developed, they are merely illustrative of the principles ofthe inventions. Other embodiments and configurations may be devisedwithout departing from the spirit of the inventions and the scope of theappended claims.

We claim:
 1. A device for measuring depth of chest compressions, saiddevice comprising: an accelerometer capable of producing an accelerationsignal corresponding to the acceleration of chest compressions; a meansfor integrating the acceleration signal, said means for integratingoperably connected to the accelerometer, and said means for integratingcapable of producing a measured depth signal corresponding to themeasured depth of chest compressions; and a processor operably connectedto the means for integrating, said processor programmed to produce anestimated actual depth signal corresponding to the estimated actualdepth of chest compressions by calculating an autoregressive movingaverage of the measured depth signal.
 2. The device of claim 1 furthercomprising a display operably connected to the processor, and whereinthe display is capable of displaying the es timated actual depth signal.3. The device of claim 1 further comprising a means for user feedbackoperably connected to the processor, wherein the means for user feedbackis capable of providing feedback that indicates whether chestcompressions fall within a pre-determined range of values.
 4. The deviceof claim 1 further comprising: a means for measuring an ECG signal of apatient, said ECG signal having a noise component and an actualcomponent; a means for estimating the noise component of the ECG signaloperably connected to the means for measuring the ECG signal and to theprocessor; wherein the processor is further programmed to produce theestimated actual depth signal when the noise component of the ECG signalfalls within a predetermined range.
 5. The device of claim 4 furthercomprising a display operably connected to the processor, and whereinthe display is capable of displaying the estimated actual depth signal.6. The device of claim 4 further comprising a means for user feedbackoperably connected to the processor, wherein the means for user feedbackis capable of providing feedback that indicates whether chestcompressions fall within a pre-determined range of values.
 7. The deviceof claim 6 wherein the processor is further capable of measuring therate of chest compressions and wherein the means for user feedback isfurther capable of displaying the rate of chest compressions.
 8. Thedevice of claim 6 wherein the processor is further capable of measuringthe duty cycle of chest compressions and wherein the means for userfeedback is further capable of displaying the duty cycle of chestcompressions.
 9. The device of claim 1 further comprising a means forperforming ventilation capable of ventilating the patient during CPR.10. The device of claim 9 wherein the means for ventilation comprises ameans for performing electro-ventilation.
 11. A device for performingchest compressions, said device comprising: a means for performing chestcompressions on a patient; an accelerometer capable of producing anacceleration signal corresponding to the acceleration of chestcompressions; a means for integrating the acceleration signal, saidmeans for integrating operably connected to the accelerometer, and saidmeans for integrating capable of producing a measured depth signalcorresponding to the measured depth of chest compressions; a processoroperably connected to the means for integrating and to the means forperforming chest compressions, said processor capable of producing anestimated actual depth signal corresponding to the estimated actualdepth of chest compressions by calculating an autoregressive movingaverage of the measured depth signal; wherein the means for performingchest compressions is capable of adjusting the depth of chestcompressions based on the estimated actual depth signal.
 12. The deviceof claim 11 wherein the means for performing chest compressionscomprises an automatic chest compression device.
 13. The device of claim11 further comprising: a means for measuring an ECG signal of a patient,said ECG signal having a noise component and an actual component; ameans for estimating the noise component of the ECG signal operablyconnected to the means for measuring the ECG signal and to theprocessor; wherein the processor is further programmed to produce theestimated actual depth signal when the noise component of the ECG signalfalls within a predetermined range.
 14. The device of claim 13 whereinthe means for performing chest compressions comprises an automatic chestcompression device.
 15. The device of claim 11 further comprising ameans for performing ventilation capable of ventilating the patientduring CPR.
 16. The device of claim 15 wherein the means for ventilationcomprises a means for performing electro-ventilation.
 17. The device ofclaim 11 further comprising a display operably connected to theprocessor, and wherein the display is capable of displaying theestimated actual depth signal.
 18. The device of claim 11 furthercomprising a means for user feedback operably connected to theprocessor, wherein the means for user feedback is capable of providingfeedback that indicates whether chest compressions fall within apre-determined range of values.
 19. The device of claim 18 wherein theprocessor is further capable of measuring the rate of chest compressionsand wherein the means for user feedback is further capable of displayingthe rate of chest compressions.
 20. The device of claim 18 wherein theprocessor is further capable of measuring the duty cycle of chestcompressions and wherein the means for user feedback is further capableof displaying the duty cycle of chest compressions.
 21. A device formeasuring depth of chest compressions, said device comprising: anaccelerometer capable of producing an acceleration signal correspondingto the acceleration of chest compressions, said acceleration signalhaving a noise component and an actual component; a processor operablyconnected to the accelerometer, said processor capable of producing,from the acceleration signal, an estimated actual acceleration signalcorresponding to the estimated actual acceleration of chestcompressions; wherein the processor further comprises: a systemidentifier operably connected to the accelerometer, said systemidentifier capable of producing an estimated noise signal correspondingto the noise component of the acceleration signal; wherein the systemidentifier produces the estimated noise signal by processing themeasured acceleration and at least one error source reference; a meansfor combining signals operably connected to the system identifier and tothe accelerometer, said means for combining signals capable of combiningthe acceleration signal and the estimated noise signal to produce theestimated actual acceleration signal; and a means for integrating theestimated actual acceleration signal, said means for integratingoperably connected to the processor, and said means for integratingcapable of integrating the estimated actual acceleration signal toproduce an estimated actual depth signal corresponding to the estimatedactual depth of chest compressions.
 22. The device of claim 21 furthercomprising a display operably connected to the processor, and whereinthe display is capable of displaying the estimated actual depth signal.23. The device of claim 21 further comprising a means for user feedbackoperably connected to the processor, wherein the means for user feedbackis capable of providing feedback that indicates whether chestcompressions fall within a pre-determined range of values.
 24. Thedevice of claim 23 wherein the processor is further capable of measuringthe rate of chest compressions and wherein the means for user feedbackis further capable of displaying the rate of chest compressions.
 25. Thedevice of claim 23 wherein the processor is further capable of measuringthe duty cycle of chest compressions and wherein the means for userfeedback is further capable of displaying the duty cycle of chestcompressions.
 26. The device of claim 21 further comprising: a means formeasuring an ECG signal of a patient, said ECG signal having a noisecomponent and an actual component; a means for estimating the noisecomponent of the ECG signal operably connected to the means formeasuring the ECG signal and to the processor; wherein the processor isfurther programmed to produce the estimated actual depth signal when thenoise component of the ECG signal falls within a predetermined range.27. The device of claim 26 further comprising a display operablyconnected to the processor, and wherein the display is capable ofdisplaying the estimated actual depth signal.
 28. The device of claim 26further comprising a means for user feedback operably connected to theprocessor, wherein the means for user feedback is capable of providingfeedback that indicates whether chest compressions fall within apre-determined range of values.
 29. The device of claim 28 wherein theprocessor is further capable of measuring the rate of chest compressionsand wherein the means for user feedback is further capable of displayingthe rate of chest compressions.
 30. The device of claim 28 wherein theprocessor is further capable of measuring the duty cycle of chestcompressions and wherein the means for user feedback is further capableof displaying the duty cycle of chest compressions.
 31. The device ofclaim 21 further comprising a means for performing ventilation capableof ventilating the patient during CPR.
 32. The device of claim 31wherein the means for ventilation comprises a means for performingelectro-ventilation.
 33. A device for performing chest compressions,said device comprising: a means for performing chest compressions on apatient; an accelerometer capable of producing an acceleration signalcorresponding to the acceleration of chest compressions, saidacceleration signal having a noise component and an actual component; aprocessor operably connected to the accelerometer, said processorcapable of producing, from the acceleration signal, an estimated actualacceleration signal corresponding to the estimated actual accelerationof chest compressions; wherein the processor further comprises: a systemidentifier operably connected to the accelerometer, said systemidentifier capable of producing an estimated noise signal correspondingto the noise component of the acceleration signal; wherein the systemidentifier produces the estimated noise signal by processing themeasured acceleration and at least one error source reference; a meansfor combining signals operably connected to the system identifier and tothe accelerometer, said means for combining signals capable of combiningthe acceleration signal and the estimated noise signal to produce theestimated actual acceleration signal; and a means for integrating theestimated actual acceleration signal, said means for integratingoperably connected to the processor, and said means for integratingcapable of integrating the estimated actual acceleration signal toproduce an estimated actual depth signal corresponding to the estimatedactual depth of chest compressions; wherein the means for performingchest compressions is capable of adjusting the depth of chestcompressions based on the estimated actual depth signal.
 34. The deviceof claim 33 further comprising a display operably connected to theprocessor, and wherein the display is capable of displaying theestimated actual depth signal.
 35. The device of claim 33 furthercomprising a means for user feedback operably connected to theprocessor, wherein the means for user feedback is capable of providingfeedback that indicates whether chest compressions fall within apre-determined range of values.
 36. The device of claim 35 wherein theprocessor is further capable of measuring the rate of chest compressionsand wherein the means for user feedback is further capable of displayingthe rate of chest compressions.
 37. The device of claim 35 wherein theprocessor is further capable of measuring the duty cycle of chestcompressions and wherein the means for user feedback is further capableof displaying the duty cycle of chest compressions.
 38. The device ofclaim 33 further comprising: a means for measuring an ECG signal of apatient, said ECG signal having a noise component and an actualcomponent; a means for estimating the noise component of the ECG signaloperably connected to the means for measuring the ECG signal and to theprocessor; wherein the processor is further programmed to produce theestimated actual depth signal when the noise component of the ECG signalfalls within a predetermined range.
 39. The device of claim 38 furthercomprising a display operably connected to the processor, and whereinthe display is capable of displaying the estimated actual depth signal.40. The device of claim 38 further comprising a means for user feedbackoperably connected to the processor, wherein the means for user feedbackis capable of providing feedback that indicates whether chestcompressions fall within a pre-determined range of values.
 41. Thedevice of claim 40 wherein the processor is further capable of measuringthe rate of chest compressions and wherein the means for user feedbackis further capable of displaying the rate of chest compressions.
 42. Thedevice of claim 40 wherein the processor is further capable of measuringthe duty cycle of chest compressions and wherein the means for userfeedback is further capable of displaying the duty cycle of chestcompressions.
 43. The device of claim 33 further comprising a means forperforming ventilation capable of ventilating the patient during CPR.44. The device of claim 33 wherein the means for ventilation comprises ameans for performing electro-ventilation.
 45. A device for estimating anactual ECG signal of a patient while performing chest compressions, saiddevice comprising: a means for performing chest compressions on apatient; a means for sensing the ECG signal of the patient, said meansfor sensing the ECG signal capable of producing a measured ECG signalcorresponding to the measured value of the ECG signal of the patient,wherein the ECG signal comprises an actual component and a noisecomponent; a compression sensor operably connected to the means forperforming chest compressions, said compression sensor capable ofproducing a compression signal corresponding to the presence of a chestcompression; a processor operably connected to the compression sensorand to the means for sensing the ECG signal, said processor capable ofproducing an estimated actual ECG signal corresponding to the estimatedactual ECG signal of the patient; wherein the processor furthercomprises: a system identifier operably connected to the compressionsensor, said system identifier capable of producing the estimated noisecomponent of the ECG signal; wherein the system identifier produces theestimated noise component of the ECG signal by processing the measuredECG signal and the compression signal; and a means for combining signalsoperably connected to the system identifier and to the means for sensingthe ECG signal, said means for combining signals capable of combiningthe measured ECG signal and the estimated noise component of the ECGsignal to produce the estimated actual ECG signal.
 46. The device ofclaim 45 wherein the compression sensor comprises a load sensor.
 47. Thedevice of claim 46 wherein the load sensor is disposed beneath thepatient.
 48. The device of claim 45 wherein the compression sensorcomprises a means for measuring the displacement of a compression belt.49. The device of claim 48 wherein the means for measuring thedisplacement of a compression belt comprises an encoder.
 50. The deviceof claim 49 wherein the encoder comprises a rotary encoder.
 51. Thedevice of claim 49 wherein the encoder comprises an optical encoder. 52.The device of claim 45 wherein the compression sensor comprises anaccelerometer.
 53. The device of claim 45 wherein the system identifiercomprises a moving average filter.
 54. The device of claim 45 whereinthe system identifier comprises an autoregressive moving average filter.55. The device of claim 45 wherein the system identifier comprises anautoregressive moving average with truncated derivative filter.
 56. Thedevice of claim 45 wherein the system identifier comprises a Kalmanfilter.
 57. The device of claim 45 wherein the system identifiercomprises a recursive least squares filter
 58. The device of claim 45wherein the system identifier comprises a recursive instrumentalvariable filter.
 59. The device of claim 45 wherein the systemidentifier comprises a recursive prediction error filter.
 60. The deviceof claim 45 wherein the system identifier comprises a recursivepseudolinear regression filter.
 61. The device of claim 45 wherein thesystem identifier comprises a recursive Kalman filter for time-varyingsystems filter.
 62. The device of claim 45 wherein the system identifiercomprises a recursive Kalman filter with parametric variation filter.63. A method of determining the estimated actual depth of chestcompressions during chest compressions, wherein the method comprises thesteps of: providing an accelerometer capable of sensing accelerationscaused by chest compressions; performing chest compressions, wherein theaccelerometer measures an acceleration signal during chest compressions;double integrating the acceleration signal to produce a measured depthsignal corresponding to the measured depth of chest compressions;calculating an autoregressive moving average of the measured compressiondepth signal; wherein the estimated actual depth of compressionscomprises the autoregressive moving average of the measured compressiondepth signal.
 64. The method of claim 63 comprising the further stepsof: providing a reference sensor capable of identifying the start of acompression; identifying the start of a compression with the referencesensor; and calculating estimated actual depth of compressions when thestart of a compression has been identified.
 65. The method of claim 64wherein the step of providing the reference sensor comprises providing asensor capable of measuring an ECG signal of a patient, and wherein themethod comprises the further steps of: measuring the ECG signal of thepatient during compressions, wherein the ECG signal comprises a noisecomponent and an actual component; identifying the noise component ofthe ECG signal; wherein the start of each compression is identified by apredetermined change in the noise component of the ECG signal.
 66. Amethod of estimating the actual depth of chest compressions during chestcompressions, wherein the method comprises the steps of: providing anaccelerometer capable of measuring the acceleration caused by chestcompressions; performing compressions, wherein the accelerometerproduces an acceleration signal corresponding to the acceleration ofchest compressions; providing the acceleration signal to a first filter,said first filter producing a filtered acceleration signal; providingthe filtered acceleration signal to a processor, said processorproducing a velocity signal corresponding to the velocity of chestcompressions upon integrating the filtered acceleration signal;providing the velocity signal to a second filter, said second filterproducing a filtered velocity signal; providing the filtered velocitysignal to the processor, said processor producing a chest depth signalcorresponding to the depth of chest compressions upon integrating thefiltered velocity signal; providing the chest depth signal to a baselinelimiter, said baseline limiter producing a chest depth signal with aprocessed baseline component; providing the chest depth signal with aprocessed baseline component to a starting point detector, said startingpoint detector identifying the starting points of chest compressions,and further providing a starting point signal corresponding to thestarting points of chest compressions to a means for combining signals;providing the chest depth signal with a processed baseline component toa peak limiter, said peak limiter producing a chest depth signal with aprocessed peak component; providing the chest depth signal with aprocessed peak component to a peak detector, said peak detectoridentifying the peaks of chest compressions corresponding to the maximumdepths of compressions, and further providing a peak signalcorresponding to the peaks of chest compressions to the means forcombining signals; calculating the estimated actual depth of chestcompressions by combining the starting point signal with the peaksignal.
 67. The method of claim 66 comprising the further step offiltering the chest depth signal with a processed baseline component.68. The method of claim 66 comprising the further step of filtering thechest depth signal with a processed peak component.
 69. The method ofclaim 66 comprising the further steps of: providing the estimated actualdepth of compressions to a means for user feedback operably connected tothe processor; and providing feedback to a user that indicates whetherchest compressions fall within a predetermined range of values.
 70. Themethod of claim 66 comprising the further steps of: providing areference sensor capable of identifying the start of a compression;identifying the start of a compression with the reference sensor; andcalculating estimated actual depth of compressions when the start of acompression has been identified.
 71. The method of claim 70 wherein thestep of providing the reference sensor comprises providing a sensorcapable of measuring an ECG of a patient, and wherein the methodcomprises the further steps of: measuring the ECG signal of the patientduring compressions, wherein the ECG signal comprises a noise componentand an actual component; identifying the noise component of the ECGsignal; wherein the start of each compression is identified by apredetermined change in the noise component of the ECG signal.
 72. Amethod of estimating the actual depth of chest compressions during chestcompressions, wherein the method comprises the steps of: providing ameans for performing chest compressions on a patient; providing a sensorcapable of measuring an ECG signal of the patient and providing anaccelerometer capable of measuring the acceleration of chestcompressions; performing chest compressions on the patient, wherein theaccelerometer produces an acceleration signal corresponding to theacceleration of chest compressions; measuring the ECG signal of thepatient, wherein the measured ECG signal comprises an actual componentand a noise component, and wherein at least part of the noise componentis caused by chest compressions; identifying the noise component of theECG signal, wherein a starting point of a compression is identified by achange in the noise component of the ECG signal; calculating theestimated actual depth of compressions by double integrating theacceleration signal when the starting point of a compression has beenidentified.
 73. A method of performing CPR on a patient, wherein themethod comprises the steps of: providing an accelerometer capable ofsensing accelerations caused by chest compressions; performing chestcompressions, wherein the accelerometer measures an acceleration signalduring chest compressions; double integrating the acceleration signal toproduce a measured depth signal corresponding to the measured depth ofchest compressions; calculating an autoregressive moving average of themeasured compression depth signal; wherein the estimated actual depth ofcompressions comprises the autoregressive moving average of the measuredcompression depth signal; providing a means for ventilating the patient;and ventilating the patient during CPR.
 74. The method of claim 73wherein the step of providing a means for ventilating the patientcomprises providing a means for electro-ventilation.
 75. The method ofclaim 73 comprising the further steps of: providing a reference sensorcapable of identifying the start of a compression; identifying the startof a compression with the reference sensor; and calculating estimatedactual depth of compressions when the start of a compression has beenidentified.
 76. The method of claim 75 wherein the step of providing thereference sensor comprises providing a sensor capable of measuring anECG signal of a patient, and wherein the method comprises the furthersteps of: measuring the ECG signal of the patient during compressions,wherein the ECG signal comprises a noise component and an actualcomponent; identifying the noise component of the ECG signal; whereinthe start of each compression is identified by a predetermined change inthe noise component of the ECG signal.
 77. A method of estimating anactual ECG signal of a patient while performing chest compressions withan automatic chest compressions device, wherein the method comprises thesteps of: providing an ECG sensor capable of measuring an ECG signal ofthe patient, said ECG sensor producing a measured ECG signal having anactual component and a noise component; providing an automatic chestcompression device disposed to provide chest compressions to thepatient, said chest compression device having a load sensor capable ofdetermining the presence of a chest compression when the load sensed bythe load sensor exceeds a pre-determined value, said load sensorproducing a compression signal corresponding the presence a chestcompression; performing compressions; providing the measured ECG signalto a system identifier; providing the compression signal to the systemidentifier; estimating the noise component of the measured ECG signalwith the system identifier by processing the measured ECG signal and thecompression signal; providing the measured ECG signal and the estimatednoise component of the measured ECG signal to a means for combiningsignals; calculating the estimated actual ECG with the means forcombining signals by combining the measured ECG signal and the noisecomponent of the measured ECG signal.
 78. The method of claim 77 whereinthe step of providing an automatic chest compression device a loadsensor comprises providing an automatic chest compression device havinga load sensor that is disposed beneath the patient during compressions.79. The method of claim 77 wherein the system identifier comprises amoving average filter.
 80. The method of claim 77 wherein the systemidentifier comprises an autoregressive moving average filter.
 81. Themethod of claim 77 wherein the system identifier comprises anautoregressive moving average with truncated derivative filter.
 82. Themethod of claim 77 wherein the system identifier comprises a Kalmanfilter.
 83. The method of claim 77 wherein the system identifiercomprises a recursive least squares filter
 84. The method of claim 77wherein the system identifier comprises a recursive instrumentalvariable filter.
 85. The method of claim 77 wherein the systemidentifier comprises a recursive prediction error filter.
 86. The methodof claim 77 wherein the system identifier comprises a recursivepseudolinear regression filter.
 87. The method of claim 77 wherein thesystem identifier comprises a recursive Kalman filter for time-varyingsystems filter.
 88. The method of claim 77 wherein the system identifiercomprises a recursive Kalman filter with parametric variation filter.89. A method of estimating an actual ECG signal of a patient whileperforming chest compressions with an automatic chest compressionsdevice, wherein the method comprises the steps of: providing an ECGsensor capable of measuring an ECG signal of the patient, said ECGsensor producing a measured ECG signal having an actual component and anoise component; providing an automatic chest compression devicedisposed to provide chest compressions to the patient, said chestcompression device having an encoder capable of determining the presenceof a chest compression, said encoder producing a compression signalcorresponding the presence a chest compression; performing compressions;providing the measured ECG signal to a system identifier; providing thecompression signal to the system identifier; estimating the noisecomponent of the measured ECG signal with the system identifier byprocessing the measured ECG signal and the compression signal; providingthe measured ECG signal and the estimated noise component of themeasured ECG signal to a means for combining signals; calculating theestimated actual ECG with the means for combining signals by combiningthe measured ECG signal and the noise component of the measured ECGsignal.
 90. The method of claim 89 wherein the step of providing anautomatic chest compression device having an encoder comprises providingan automatic chest compression device having an optical encoder.
 91. Themethod of claim 89 wherein the step of providing an automatic chestcompression device having an encoder comprises providing an automaticchest compression device having a rotary encoder.
 92. The method ofclaim 89 wherein the system identifier comprises a moving averagefilter.
 93. The method of claim 89 wherein the system identifiercomprises an autoregressive moving average filter.
 94. The method ofclaim 89 wherein the system identifier comprises an autoregressivemoving average with truncated derivative filter.
 95. The method of claim89 wherein the system identifier comprises a Kalman filter.
 96. Themethod of claim 89 wherein the system identifier comprises a recursiveleast squares filter
 97. The method of claim 89 wherein the systemidentifier comprises a recursive instrumental variable filter.
 98. Themethod of claim 89 wherein the system identifier comprises a recursiveprediction error filter.
 99. The method of claim 89 wherein the systemidentifier comprises a recursive pseudolinear regression filter. 100.The method of claim 89 wherein the system identifier comprises arecursive Kalman filter for time-varying systems filter.
 101. The methodof claim 89 wherein the system identifier comprises a recursive Kalmanfilter with parametric variation filter.
 102. A method of estimating anactual ECG signal of a patient while performing chest compressions withan automatic chest compressions device, wherein the method comprises thesteps of: providing an ECG sensor capable of measuring an ECG signal ofthe patient, said ECG sensor producing a measured ECG signal having anactual component and a noise component; providing an automatic chestcompression device disposed to provide chest compressions to thepatient, said chest compression device having an accelerometer capableof determining the presence of a chest compression, said accelerometerproducing a compression signal corresponding the presence a chestcompression; performing compressions; providing the measured ECG signalto a system identifier; providing the compression signal to the systemidentifier; estimating the noise component of the measured ECG signalwith the system identifier by processing the measured ECG signal and thecompression signal; providing the measured ECG signal and the estimatednoise component of the measured ECG signal to a means for combiningsignals; calculating the estimated actual ECG with the means forcombining signals by combining the measured ECG signal and the noisecomponent of the measured ECG signal.
 103. The method of claim 102wherein the system identifier comprises a moving average filter. 104.The method of claim 102 wherein the system identifier comprises anautoregressive moving average filter.
 105. The method of claim 102wherein the system identifier comprises an autoregressive moving averagewith truncated derivative filter.
 106. The method of claim 102 whereinthe system identifier comprises a Kalman filter.
 107. The method ofclaim 102 wherein the system identifier comprises a recursive leastsquares filter
 108. The method of claim 102 wherein the systemidentifier comprises a recursive instrumental variable filter.
 109. Themethod of claim 102 wherein the system identifier comprises a recursiveprediction error filter.
 110. The method of claim 102 wherein the systemidentifier comprises a recursive pseudolinear regression filter. 111.The method of claim 102 wherein the system identifier comprises arecursive Kalman filter for time-varying systems filter.
 112. The methodof claim 102 wherein the system identifier comprises a recursive Kalmanfilter with parametric variation filter.